Mechanisms and methods for the design and fabrication of a mechanical interface between a wearable device and a human body segment

ABSTRACT

The system includes an instrument for determining the anatomical, biomechanical, and physiological properties of a body segment that includes one or more force sensitive probes is provided. A human operator actuates one or more force sensitive probes, wherein the force sensitive probes are positioned at the surface of the body segment. The operator pushes on the force sensitive probes with varying force applied on the body segment to measure tissue deflection forces. The instrument may include one or more of gyroscopes, accelerometers, and magnetometers capable of measuring changes in tissue deflection caused by the force sensitive probes relative to a grounded reference frame in 3-D space, wherein the tissue deflection force data and the change in tissue deflection data are used to compute segment tissue viscoelastic properties. The instrument may also be untethered or wireless.

CROSS REFERENCE TO RELATED APPLICATION

This application is related to and claims priority from earlier filedprovisional patent application Ser. No. 62/043,842, filed Aug. 29, 2014,the entire contents thereof is incorporated herein by reference.

BACKGROUND OF THE INVENTION

This invention relates generally to the design and fabrication of amechanical interface that connects a human body segment to wearabletechnology such as shoes, bras, apparel, seats, limb prostheses, limborthoses or body exoskeletal devices. Conventional design andfabrication strategies for a mechanical interface employ an incompletedata representation of the relevant human body segment, anon-quantitative methodology to determine the corresponding interfacedesign, and inadequate fabrication techniques to construct the finalproduct.

To design and fabricate a socket for a prosthetic limb, for example, aprosthetist first takes a mold of the residual limb, capturing its 3-Dshape. Depending on the practitioner's preference, this molding processis performed when the relevant human body segment is either in a loadedor unloaded state. The measurement of residual limb shape is mosttypically performed using a plaster-impregnated gauze that is firstdipped into water and then wrapped around the residual limb. Oncewrapped, the plaster hardens to form a female cup that is then pouredwith plaster to form a male plug with the residual limb's shape. Theprosthetist then removes plaster in soft tissue regions where he/shewants the final socket interface to compress the residual limb tissue,and adds plaster around sensitive regions to create a void in the finalsocket wall. Once these craft modifications are complete, a final carboncomposite or thermoplastic socket is fabricated over the male plug. Asan example, in FIGS. 1A and 1B, different views of a plaster male plug30 of a below-knee residual limb is shown after a prosthetist hadcompleted modifications to the unloaded shape, and in FIGS. 2A and 2B,different views of the final, carbon-fiber transtibial socket 32 isshown, fabricated using the plaster male plug 30 shown in FIGS. 1A and1B. The residual biological limb is then inserted into the carbonsocket, with a often silicone liner and socks worn on the residual limbto cushion the limb and to minimize skin-interface chaffing.

More specifically, FIGS. 1A and 1B, for a transtibial patient, show aplaster male plug 30 after regions of soft tissue had been reduced fromthe original residual-limb shape, and sensitive regions expanded upon.

In FIG. 2A, an anterior view of the final carbon transtibial socket 32is shown while FIG. 2B shows an internal view of the same socket 32,depicting internal craft modifications performed by a prosthetist tooptimize comfort. To construct the socket 32, a prosthetist fabricatescarbon materials over the male plug shown in FIGS. 1A and 1B.

It is well known in the art, as this particular example illustrates,today's design and fabrication strategies for mechanical interfacesemploy an incomplete data representation of the relevant human bodysegment, and a non-quantitative methodology to determine thecorresponding interface design. Furthermore, today's interfacefabrication strategies do not allow for continuously varying materialproperties within the interface that reflect the multi-tissue,continuously-varying, viscoelastic properties of the underlying anatomyfor which the mechanical interface is designed to intimately connect.Such a poor correspondence between body and synthetic interface causesdiscomfort for the wearer due to excessive pressures, internal strains,shear forces and skin chaffing between the attached device, clothing orshoe article, and the human body segment.

As noted earlier, practitioners typically measure the shape of the humanlimb segment of interest, and then modify that shape usingnon-quantitative craft techniques that do not quantitatively map theunderlying anatomical, biomechanical and physiological features totissue compression levels and internal stresses and strains imposed bythe interface. Moreover, the final interface is typically homogenious,or nearly homogenious, in terms of its viscoelastic properties,spatially and temporally; for example, the carbon fiber socket shown inFIG. 2 is rigid across the entire interface surface, and that rigidityis invariant in time. Further, the silicone liner worn directly againstthe residual skin is also typically homogeneous, or nearly so, in termsof its viscoelastic properties.

Attempts have been made to vary the viscoelastic properties of theinterface spatially using a ‘windowing’ approach where holes are cutinto a rigid, external interface wall to allow an intermediate, softermaterial to penetrate through the window upon load bearing applied tothe interface. However, such windowing techniques use separate distinctmaterial components resulting in an interface that does not reflect thecontinuously-varying human body viscoelastic properties found in theunderlying anatomy. Further, often the tensile elasticity of thesilicone liner, worn on the residual limb in the case of leg amputation,is varied somewhat spatially so as to stiffen the liner against axial,longitudinal stretch, but to still allow compliance for circumferentialtensile strains. However, these liner impedance variations do notreflect the multi-tissue, continuously-varying, viscoelastic propertiesof the underlying anatomy.

In view of the foregoing, there is a demand for a system that can moreeffectively and accurately determine the anatomical, biomechanical, andphysiological properties of a body segment in order to provide asuperior mechanical interface between a wearable device and human bodysegment.

SUMMARY OF THE INVENTION

The present invention preserves the advantages of prior art mechanismsand methods for the design and fabrication of a mechanical interfacebetween a wearable device and a human body segment. In addition, itprovides new advantages not found in currently available mechanisms andmethods and overcomes many disadvantages of such currently availablemechanisms and methods.

The invention is generally directed to an instrument, preferablyuntethered, for determining the anatomical, biomechanical, andphysiological properties of a body segment that includes one or moreforce sensitive probes, a human operator to actuate the one or moreforce sensitive probes, wherein the one or more force sensitive probesare positioned at the surface of the body segment and the operator thenpushes on the one or more force sensitive probes with varying forceapplied on the body segment to measure tissue deflection forces, whereinthe untethered instrument further comprises one or more of gyroscopes,accelerometers, and magnetometers capable of measuring changes in tissuedeflection caused by the one or more force sensitive probes relative toa grounded reference frame in 3-D space, wherein the tissue deflectionforce data and the change in tissue deflection data are used to computesegment tissue viscoelastic properties.

It is therefore an object of the present invention to provide a systemincludes an instrument for determining the anatomical, biomechanical,and physiological properties of a body segment that includes one or moreforce sensitive probes is provided.

A further object of the present invention is to enable a human operatorto actuates one or more force sensitive probes, wherein the forcesensitive probes are positioned at the surface of the body segment wherethe operator pushes on the force sensitive probes with varying forceapplied on the body segment to measure tissue deflection forces.

Yet another object of the present invention is to provide an instrumentmay include one or more of gyroscopes, accelerometers, and magnetometerscapable of measuring changes in tissue deflection caused by the forcesensitive probes relative to a grounded reference frame in 3-D space,wherein the tissue deflection force data and the change in tissuedeflection data are used to compute segment tissue viscoelasticproperties.

Yet a further object of the present invention is to provide aninstrument that is untethered or wireless.

BRIEF DESCRIPTION OF THE DRAWINGS

The patent or application file contains at least one drawing executed incolor. Copies of this patent or patent application publication withcolor drawing(s) will be provided by the Patent and Trademark Officeupon request and payment of the necessary fee.

The novel features which are characteristic of the present invention areset forth in the appended claims. However, the invention's preferredembodiments, together with further objects and attendant advantages,will be best understood by reference to the following detaileddescription taken in connection with the accompanying drawings in which:

FIGS. 1A and 1B show, for a transtibial patient, different views of aplaster male plug is shown after regions of soft tissue had been reducedfrom the original residual-limb shape, and sensitive regions expandedupon;

FIG. 2A shows and external anterior view of the final carbon transtibialsocket;

FIG. 2B shows an internal view of the socket of FIG. 2A;

FIGS. 3A-C show three poses of a transtibial residual limb correspondingto a particular knee flexion angle;

FIGS. 4A-C show the coordinate information from three triangulated posesof a transtibial residual limb of FIG. 3AC that are used to compute thestrain transforms;

FIG. 5A shows the average strain of each triangular face being analyzedand mapped to a color, where the skin strain levels are shown for thepartially flexed pose;

FIG. 5B shows the average strain of each triangular face being analyzedand mapped to a color, wherein skin strain levels are shown for thefully flexed pose;

FIG. 6A is a strain field of the knee flexed to approximately 90degrees;

FIG. 6B is the detail A indicated in FIG. 6A;

FIG. 6C is the detail B indicated in FIG. 6A;

FIG. 7A shows a 3D view of bones and patella tendon is shown for theright residual limb of a transtibial amputee;

FIG. 7B shows the orthogonal distance D between the unloaded skinsurface and a bone intersection;

FIG. 7C shows another view of the orthogonal distance D between theunloaded skin surface and a bone intersection;

FIGS. 8A and 8B show different views of a single probe on a flexiblearm;

FIG. 9 shows a single probe on a flexible arm collecting anatomical,biomechanical and physiological data of a transtibial residual limb;

FIGS. 10A and 10B show different views of a probe array on a flexiblearm;

FIG. 11 shows two probe arrays on flexible arms collecting anatomical,biomechanical and physiological data of a transtibial residual limb;

FIGS. 12A and 12B different views of a finger probe on a flexible arm;

FIG. 13 shows finger probes on flexible arms collecting anatomical,biomechanical and physiological data of a transtibial residual limb;

FIG. 14A shows an untethered finger probe;

FIG. 14B shows a cross-sectional view of the finger probe of FIG. 14A;

FIGS. 15A and 15B show different views of multiple untethered fingerprobes;

FIGS. 16A and 16B show multiple untethered finger probes collectinganatomical, biomechanical and physiological data of the ankle-footcomplex;

FIG. 17 is a graph showing linear and non-linear relationships betweenthe body's viscoelastic properties as estimated from soft tissue depthplotted horizontally, and the corresponding durometer of the mechanicalinterface plotted vertically;

FIG. 18 is a graph showing relationships between the unloaded interfaceshape and tissue stiffness approximated here as soft tissue depth;

FIG. 19 is a perspective view of a 3D printer;

FIGS. 20A-20D show MRI images for the right leg of a transtibialamputee;

FIGS. 20E-20H show images of the soft tissue depth model of the residuallimb;

FIGS. 20I-20L show a 3-D printed prosthetic socket where every materialcolor corresponds to a material having a distinct durometer and tensilestrength;

FIGS. 20M-20P show the socket's most rigid, high tensile strengthmaterial being is modeled;

FIG. 21 shows The Von Mises Stress distribution for finite elementanalyses shown in the fourth row shown in FIGS. 20M-20P;

FIG. 22 is a graph mapping between the Young's Modulus of socketinterface materials shown in the third row of FIGS. 20I-20L to the softtissue depth at each location shown in the second row of FIGS. 20E-20H,which is color coded by categories of soft tissue depth;

FIG. 23 is a transtibial socket design;

FIG. 24 shows a thin compliant material, or liner, bonded at its distalaspect to the multi-material prosthetic socket shown in FIG. 23 to forma fully-integrated mechanical interface with the body;

FIG. 25 is a full view of the multi-material prosthetic socket with theinternal liner and socket bonded within and a carbon fiber outermaterial;

FIG. 26 shows the internal liner bonded to the inner surface of thevariable-impedance socket within a carbon fiber element;

FIG. 27 is shows the outer carbon fiber element designed for structuralintegrity;

FIG. 28 shows a front view of the liner embedded with sensingmodalities;

FIG. 29 shows an electro-active polymer electrode;

FIG. 30A shows the average strain of each triangular face being analyzedand mapped to a color with skin strain levels shown for the partiallyflexed pose.

FIG. 30B shows the average strain of each triangular face being analyzedand mapped to a color with skin strain levels shown for the fully flexedpose; and

FIG. 31 is a table (Table 1) showing the color mapping used in FIG. 15.

DETAILED DESCRIPTION OF THE INVENTION

In accordance with the present invention, as a resolution to thedifficulties discussed above, a quantitative methodology is presentedthat relates human-body anatomical, biomechanical and physiologicalproperties to the design and fabrication of a novel mechanicalinterface. Specifically, the present invention describes a quantitative,scientific methodology that relates measurements of biological segmentshape, skin strain characteristics resulting from body movement,viscoelastic tissue properties for state disturbances perpendicular tothe bodies surface, sensitivities to applied pressure due to bursitis,nerve, blood flow restrictions, chronic wounds, etc., vascularizationand peripheral nerve anatomy, and the like, to the correspondinginterface shape and impedance characteristics, spatially and temporally.It will be understood by those of ordinary skill in the art that themethodologies presented in herein could be employed in themechanical-interface design and fabrication of any wearable article ormechanism, including shoes, cloths, seats, bras, prostheses, orthoses orexoskeletons.

The design and fabrication methodologies of the present invention aredivided into three different phases or steps. The first step comprisesacquiring a comprehensive data set of the relevant human body segment'sunderlying anatomy, biomechanics and physiology, and then processingthese data to build a digital representation, or model, of thebiological segment for which the mechanical interface will connect.Next, in a second step, a quantitative mapping from the biological modelto an interface model is generated. The interface model is a digitalrepresentation of interface shape and impedance properties. Finally, ina third step, the interface model is used to fabricate either a testinterface, or the final interface to be used by the wearer of thearticle or mechanism.

Step 1: Biological Data Acquisition and Modeling

In accordance with the present invention, the first part to theproduction of a mechanical interface includes collecting anatomical,biomechanical and physiological data that can be used to develop a modelof the biological segment of interest. Such a model is necessary todescribe the relevant biological segment's properties, including but notlimited to its shape, skin strain characteristics resulting from bodymovement, viscoelastic tissue properties for state disturbancesperpendicular to the bodies surface, sensitivities to applied pressuredue to bursitis, nerve, blood flow restrictions, chronic wounds, and thelike, all as a function of anatomical location. Such a data-driven modelcan be represented as a vector of biological properties at eachanatomical location across the body segment for which an article ofclothing, a worn shoe, or a wearable device is designed to interface.

Data Types and Methods of Acquisition

Skin Strain

A critical data set relevant to the design of a mechanical interface isskin strain dynamics caused by body joint movements. A procedure isoutlined in this section that can be used to collect data necessary toestimate the skin strain field of the biological segment of interest,and then to compute the skin strain field as a function of limb posture.Such information is necessary to understand how the mechanical interfaceshould move and stretch relative to the skin surface, so as to minimizeshear forces and discomfort at the skin-interface junction.

In this procedure, the biological limb is first marked with a matrix ofsmall (˜2 mm diameter), black-ink dots across the entire skin-surfacearea for which the interface is designed to interact. The specificanatomical location and distance between these dots need not be precise,but the resolution, or the number of dots per cm² is important, as thisresolution defines the resolution of the resulting skin strain field.Further, the pattern of dots should be randomized, providing theopportunity to create a unique skin speckle pattern for each anatomicalregion. With such a patterning, a single camera image taken of a smallregion of the skin surface can be used to determine the anatomicalposition at which the camera's lens is pointed or directed. Such ananatomical-positioning algorithm can be achieved by comparing thesingle, anatomically-local image to the full speckle patterns across theentire biological segment. As an alternative to a matrix of small dots,the skin of the biological segment of interest can be speckled with asponge where the sponge is first dipped in FDA approved body paint. Bydabbing the painted sponge across the skin surface, a unique pattern ofskin speckles can be created.

Next, separate poses, or joint postures of the biological segment ofinterest, are captured using photogrammetric tools. Using approximately30 digital photographs for each limb pose, 3D models can be generated.

The coordinates of the black dots on the skin are marked and exportedfor analysis. The point clouds for each pose are triangulated in acorresponding manner so the mapping of points to triangles is the same.In FIG. 3, an example is shown for a transtibial amputee's residual limbshowing three levels of knee flexion, and a matrix of black dots acrossthe skin surface.

Referring now to FIGS. 3A-C, three different poses of a transtibialresidual limb 34 are shown each corresponding to a particular kneeflexion angle. Black dots 110 mark the skin at a resolution ofapproximately 4 dots per cm².

The black dots 110 are the nodes of the finite element model and serveas the vertices for the surface triangulation. FIGS. 4A-C show therespective triangulated surface corresponding to the poses displayed inFIGS. 3A-C.

Further, FIGS. 4A-C show the coordinate information from threetriangulated poses of a transtibial residual limb are used to computethe strain transforms. A constant strain element analysis is performedon each triangle to ascertain the strain field of the limb's surface.

The deformation of each triangular element from one pose to another isdecomposed into a translation, rotation, and stretch via an affinetransform. The three initial coordinate pairs (x_(i), y_(i)) and threefinal coordinate pairs (x_(f), y_(f)) are used to find the affinetransform linking the two limb poses. Equation 1 represents the affinetransformation matrix that links the point sets for each element. Therigid body translation from the initial to the final pose (Δx, Δy) isneglected as it has no effect on the strain within the element.

$\begin{matrix}{\begin{bmatrix}x_{f} \\y_{f} \\1\end{bmatrix} = {\begin{bmatrix}\; & \; & \; & {\Delta\; x} \\\; & A & \; & \; \\\; & \; & \; & {\Delta\; y} \\0 & \; & 0 & 1\end{bmatrix}\begin{bmatrix}x_{i} \\y_{i} \\1\end{bmatrix}}} & (1)\end{matrix}$

The matrix A is a 2×2 matrix that contains the information about how aparticular triangle is rotated and stretched. A singular valuedecomposition (SVD) of matrix A isolates the components of thedeformation as described by equation 2. The SVD interprets thetransformation as a rotation V* to the principal coordinate frame,followed by a stretch Σ along those axes, and an additional rotation Uto the final coordinate frame.A=UΣV*  (2)

The stretch matrix yields the principal strains which are used tocompute the average strain of each constant strain triangle. Equation 3computes the von Mises or equivalent strain ∈_(e) from the principalstrains, ∈₁ and ∈₂. FIG. 5 shows the equivalent strain of eachtriangulation resulting from the deformation of the original, extendedpose to two different levels of knee flexion. The average strain is ascalar value that is useful for assessing the overall stretch of anelement.

$\begin{matrix}{ɛ_{e} = {\frac{1}{2}\sqrt{\left( {ɛ_{1} - ɛ_{2}} \right)^{2} + ɛ_{1}^{2} + ɛ_{2}^{2}}}} & (3)\end{matrix}$

Furthermore, the strain state of each two-dimensional surface elementcan be derived from Mohr's circle using the principal straininformation. This maps the two principal strains to a combination ofnormal and shear strains in another coordinate frame.

The average strain of each triangular face is analyzed and mapped to acolor. Skin strain levels are shown for the partially flexed pose, asshown in the plot of FIG. 5A, and the fully flexed pose, as shown in theplot of FIG. 5B. Here higher average strain is shown around the kneepatella due to the right pose's increased knee flexion.

The strain field can be computed using the information from the SVD ofeach triangle. FIG. 6A shows a plot of the strain field for theparticular case of a transtibial amputation. The red vectors representthe direction and magnitude of the larger of the two normal strains ofeach triangle. The blue vectors represent the smaller strain. Any shearstrain is represented by the angle between the corresponding red andblue vectors of a particular triangular element. The strains throughouteach triangle are assume to be constant and are therefore plotted at thecentroid of each triangle. If a high enough dot resolution is used, aconstant strain element analysis is sufficient to assess the strainstate of a deformed surface.

More specifically, FIG. 6 shows the strain field of the knee flexed toapproximately 90 degrees. It should be noted how the larger (red) strainfield is nearly horizontal proximal to the knee joint as the skinstretches circumferentially to accommodate the contractions of thehamstring and quadriceps muscles, as shown in Detail A of FIG. 6B.Further, large strains are highly longitudinal at the knee patella, andjust proximal to the patella, as the skin stretches over the patelladuring knee flexion, as shown in Detail B of FIG. 6C. Here, thelongitudinal direction is along the long axis of the femur.

Segment Shape

A critical data set relevant to the design of a mechanical interface isunloaded body shape. Using the photogrammetric data collection procedureoutlined in the previous section, the unloaded shape of the biologicalsegment at any given static joint pose can be ascertained simply byfitting a shape model to the coordinate x, y, z data of each skin dot ormarking, such as in FIGS. 4A-C. Other non-contact instruments can alsobe used to measure the shape of a biological segment, using for exampleimaging tools such as three-dimensional laser scanners, MRI orultrasound. Alternatively, instruments designed to be in physicalcontact with the biological body segment of interest can be used toestimate its unloaded shape. Using an instrument with skin contact andposition sensing, the x, y, z location of any point on the skin surfacecan be measured by simply touching the skin lightly at that point. Byrepeatedly touching the skin surface at a high resolution of skinpoints, the shape of the limb can ultimately be determined.

Tissue Viscoelastic and Sensitivity Properties

Other critical data relevant to the design of a mechanical interface aretissue viscoelastic and sensitivity properties for orthogonal bodycompressions. Instruments can be used to estimate the impedance(stiffness and damping) of body tissue through physical contact with thebiological body. Here the impedance is measured directly by theinstrument by physically applying a force to the body at a single point,or at several points across a region, or through the application of apressure across a broader region. As force is applied perpendicular tothe body's surface at each anatomical point, or node of interest, theinstrument's sensors can measure the global three-dimensional locationof the body surface point where the force is being applied, the appliedtissue deflection, time rate of change of tissue deflection and theapplied force or pressure. These measured data can then be used toestimate tissue impedance, or viscoelastic property, of the body segmentas a function of anatomical location, thereby establishing aviscoelastic map of the body part or segment.

In addition to these viscoelastic tissue measurements, by way of theinstrument's physical contact with the body surface, a directquantitative measurement of sensitivities to applied force/pressure canbe ascertained. For example, relevant to shoe design, the foot-anklecomplex may have sensitive areas caused by a bunion deformity, Haglund'sdeformity or Heel Spur Syndrome all having the potential to lead to thedevelopment of painful bursitis. A shoe's design should take intoaccount such sensitivities by using not only a tissue viscoelastic mapbut also a Sensitivity Map. A sensitivity map comprises the tissuestress/strain threshold at each anatomical point that first results indiscomfort for the subject at a maintained stress/strain application.When such an instrument in physical contact with the subject's bodyexerts a force/pressure at a particular anatomical location, as theapplied force/pressure increases, there will be a point when the subjectfirst experiences discomfort at a particular stress or strain level oftissue deformation. A sensitivity map shows this critical level, orthreshold, of stress, or strain, as a function of anatomical location.In the design of the mechanical interface, load can then be mitigatedfrom key anatomical areas of sensitivity to reduce internalstrains/stresses and wearer discomfort.

Blood Flow and Nerve/Spinal Conduction Dynamics

Imaging tools such as ultrasound can be used to measure blood flowdynamics and nerve/spinal conduction. Such data correlated to anatomicallocation are critical to the design of a wearable garment or device,since external loads applied to the biological segment from the wearablemight alter such dynamics and cause a health problem and/or discomfortfor the wearer.

Ultrasound is most useful for observing soft tissue structure within thebody. In fact, hard tissue degrades the quality of ultrasound images andimpedes the visibility of soft tissue behind/beneath it. Ultrasound ischaracterized by its sound frequency, ranging between 2-15 MHz, which ismuch higher than the audible range of humans (20-20,000 Hz) but lowenough to not seriously agitate living tissue. It also contains noionization radiation, making it safer in higher doses than x-rayimaging.

The key component of ultrasound devices is the transducer. It is thepiece responsible for sending and receiving sound waves, controlled bythe piezoelectric effect. When an electric signal is applied to acrystal within the transducer, it emits a sound at a given frequency.This is the piezoelectric effect. The emitted sound travels through thebody and is reflected by tissues. Since the piezoelectric effect is areversible process, reflected sound waves will yield an electricalsignal from the transducer's crystal when they interact with itfollowing reflection. A computer interprets these signals using opacityto symbolize tissue density. A variety of transducers exist with sizes,shapes, and functions that are designed to make them more useful forspecific tasks.

Image generation is dictated by the rate at which sound waves arereflected within the body. The speed of sound within the human body isknown to be 1540 meters per second, which makes possible the calculationof tissue depth given reflection time (i.e. time of signal return). Maxdepth and resolution are a function of frequency. Higher frequencieshave better resolution but do not penetrate as deeply. Higherfrequencies equal higher attenuation. Generally, it is best to choosethe highest possible frequency that will achieve the depth of interest.To determine the achievable depth of a specific sound frequency, use thefollowing equation:

$\begin{matrix}{d_{\max} = \frac{v_{sb}}{f}} & (4)\end{matrix}$where d_(max) is the maximum achievable depth, v_(sb) is the speed ofsound within the human body, and f is the sound frequency. Typicalfrequencies for deep body imaging range from 1.5 to 3 MHz, whilefrequencies for superficial structures range from 5 to 10 MHz.Transducers are often characterized by their ability to yield a range offrequencies and must be chosen accordingly.

Biological-Limb Modeling

After the biological limb is captured using photogrammetric tools, thebiological limb of interest can be imaged with a MRI machine and/or anelectromechanical instrument can be used for measuring biological-limb,viscoelastic and sensitivity properties. Once these additional data arecollected, a grid of resolution matched to the skin of the patient (e.g.average 1×1 cm) is established where a node of known variables iscreated around each grid or averaged for a defined grid. Alternatively,the grid could correspond to the grid of skin-strain triangles, forwhich FIGS. 6A-C provide an example, where a node is the center pointwithin each respective triangle. Each biological node vector V(i) hasproperties including, but not limited to, anatomical 3D location with notissue load, maximal skin tensile strain due to joint movement,orthogonal compression stiffness K and damping B as a function of tissuecompression and compression rate, and the sensitivity toexternally-applied pressure influenced by blood flow and nerveconduction dynamics, and the presence of chronic tissue wounds such asbursitis. Here the compression stiffness and damping, or impedance, isdefined as the biological limb's response to a displacement impulseperpendicular to the skin at each node. Further, the maximum skintensile strain is computed as the average strain of the three legs ofthe corresponding strain triangle (See FIG. 5 as an example).

As an example, FIGS. 7A-C show a simple model of the residual limb of atranstibial amputee generated from MRI data. The model provides theunloaded shape and soft tissue depth of the residual limb as a functionof anatomical location. Here soft tissue depth, D, is defined as theperpendicular distance from a node skin surface area and theintersection of that line with a bone. Although tissue depth correlatesapproximately to body stiffness, K, it is understood that a moresophisticated modeling exercise of soft tissue biomechanics wouldproduce a more precise model of the residual limb's compressionstiffness, K, and damping, B, properties as a function of anatomicallocation and neural activation. Here neural activation is included sincelarge changes in viscoelastic properties occur depending upon whethermuscles are activated or relaxed. Such a biological segment model wouldalso include information on the locations of nerve and veins, and theirrelative pressure tolerances.

In FIG. 7A, a 3D view of bones and patella tendon is shown for the rightresidual limb of a transtibial amputee. The images in FIGS. 7B and 7Cshow the orthogonal distance D between the unloaded skin surface and abone intersection. Here, red regions show large tissue depths, yellowregions moderate depths, and green regions small depths. For these depthmodels, the patella tendon was removed, exposing the soft tissue depthin the region of the patella tendon just distal to the patella (shown asthe red region in the image of FIG. 7B).

Data Acquisition Instruments

Synthetic-Actuated Instruments

Instruments in physical contact with the subject's body comprisingsynthetic actuator(s) can be used to estimate 1) the orthogonalimpedance of body tissue at each anatomical point (viscoelastic map),and 2) the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point (sensitivity map). Hereorthogonal impedance refers to tissue stiffness and damping propertiesfor a tissue state (position and velocity) disturbance directedorthogonal to the surface of the skin at each anatomical location acrossthe body segment of interest.

A viscoelastic map can be ascertained through a three-step process.First, the tissue is measured by sensors where actuators apply a seriesof controlled interactions that deflect the tissue. Second, thedata—position and force with respect to time—is conditioned for systemidentification purposes. Lastly, the data are used to identify a linearor non-linear transfer function which describes the physical response ofthe tissue to a given load (force) or deflection.

The collected data consist of positions and forces that are referencedto time. This time reference allows velocity and acceleration to becalculated as well. In order to identify the system, we will look at theinput versus the output of the system. Let's say the input is X(t) andthe output is Y(t). In order to get an idea of the linear transferfunction, we first take the autocorrelation of the input function X(t)to get X_(ac)(t). We then take the cross correlation of the input andoutput to get XY_(cc)(t). Next a specialized matrix called the Toeplitmatrix is formed with X_(ac)(t): TP(t). Then, the impulse responsefunction of the system, h, is F_(s)(TP(t)⁻¹·XY_(cc)(t)). Where F_(s) isthe frequency of the samples and TP(t)⁻¹ is the inverted Toeplitzmatrix. Given a linear system, the parameters of the transfer functioncan be determined from the impulse response.

Impedance data can be collected using a ring of linear actuators thatsurround the biological limb to be mapped. Such an actuator ring iscapable of measuring every point on the ring at the same time. Between 1and 50 points (or as many as space allows) can be measuredsimultaneously with this method. Each linear actuator must beindependently controlled with its own force and position sensors. Asimpler device could be used comprising a single actuator butconsiderably more time would be required to measure tissue impedance athigh resolution across the biological segment of interest.

Human-Actuated Instruments

Alternatively, as will be described herein, a human-actuated probe orprobes, can be used to map the anatomical, biomechanical andphysiological properties of a body part for which a wearable device isto interface. Another way of measuring the body's orthogonal impedanceat each anatomical point is with a location aware instrument that has aforce sensitive probe, or force sensitive probes. With such aninstrument, a force sensing probe or probes is pushed against thesubject's body part of interest where the three-dimensional position ofthe tip of the probe(s) is measured by the instrument at all times.Additionally, if the body part under measurement is not stationary, theinstrument must also track its location in the same three-dimensionalreference frame as the measurement probe, or probes. Such a probe, orprobes, can be positioned at the surface of the human body in aperpendicular orientation to the surface area of the skin at eachanatomical point, and the probe user (practitioner/clinician/user) canthen push with varying force, compressing the subject's tissues. Sincethe force-sensing probe, or probes, measure(s) position, both of itselfand of the biological part, both the viscoelastic and sensitivity mapscan be ascertained for the body part of interest.

The probe can measure force with a simple spring and linear positionmeasurement device. Through a measurement of the deflection of thesensor's physical spring, the force can be estimated using theforce-deflection relation of the spring (e.g. F=−kx). The probe couldalso have a force sensor that is either capacitive, resistive,piezoelectric based, strain-gauge based, or any other force sensingtechnology. In addition, the probe can also include ultrasound to imagethe body to ascertain internal tissue properties and blood flow andnervous tissue transduction dynamics, and how such dynamics change asincreasing force is applied on the tissue. Ultrasonic transducers on theprobe's tip can be used to gather very detailed tissue density data,soft tissue depth (orthogonal distance from the bodies surface to thebone), and blood flow dynamics (e.g. how blood flow is altered uponincreasing applied external pressure).

The position sensing system can be physically connected to the probe,such as a structure of linkages similar to that of an industrial robotarm where each linkage has angular position sensors that are capable ofdetermining the exact position of the probe's tip from a groundedreference frame. In addition, the pen could have markers on its surfacethat could be seen by cameras. Such cameras can be used to triangulatethe position of the probe or probes. The probe could also broadcastelectromagnetic signals that are picked up by nearby electromagneticsensors for the purposes of determining the position of the probe.Furthermore, the probe could use a combination of gyroscopes,accelerometers, and magnetometers to aid in determining its position in3-D space.

The photogrammetric, force, position, velocity, acceleration andultrasound data from the probe can be communicated wirelessly or wired.The wireless method can be IR-based, Bluetooth, or any other wirelesscommunication method such as an open electromagnetic frequency.

The location of the human body part under measurement can be determinedin much the same way as the location of the probe. Using electromagneticsignals, accelerometers, gyroscopes, magnetometers, passive locationmarkers located on the biological limb and external lab frame camerasthat measure the locations of these markers, active markers on thebiological limb and receivers positioned off the limb in lab frame, orany other location technology or combination of location technologies.

Hybrid-Actuated Instruments

Alternatively, as will be described herein, a third category ofinstrument comprises both synthetic actuation and human-poweredactuation.

In accordance with the present invention, four embodiments will bedescribed that fall within the instrument categories of human-actuatedand hybrid-actuated. Each embodiment's design, and its advantages anddisadvantages are described herein.

Embodiment I Single Probe on a Flexible Arm

A position-aware, force and ultrasound probe can be used to collectanatomical, biomechanical and physiological data describing a biologicalsegment of interest. The single probe embodiment is human-actuated sincemuscle action from the user of the probe is used to apply probepositions and forces around, and onto, the biological segment. Thesingle probe on a flexible arm is shown in FIG. 8A with a close up viewshown in FIG. 8B.

The single probe 96 is attached to a flexible arm 97 from a base 98attached to a stationary lab frame location. The arm is flexible withrotary joints, e.g. at 99, in order to orient the probe tip to anylocation within an extensive 3-D volume. Precision encoders 100 arelocated throughout the flexible arm 97 to allow for real time estimatesof the location of the probe tip 102 in 3-D space.

The probe itself comprises photogrammetric, kinetic, and ultrasoundsensing. Four small cameras 101 are positioned around the longitudinalaxis of the probe. In addition, male probe tip 102 moves linearly intoand out of female probe housing 104 when forces are applied to probe tip102. A compression spring 106 and a linear potentiometer 108 serve asthe force sensor within female probe housing 104. When probe tip 102 ispushed onto a biological body segment, compressing its tissue, a forceis exerted on force sensor spring 106. Force sensor spring 106compresses against spring block 107 mechanically grounded to femaleprobe housing 104. The compression of sensor spring 106 is then measuredby linear potentiometer 108. Since the stiffness of sensor spring 106 isknown, the sensing of spring compression provides force information. Theprobe also comprises ultrasound. Within probe tip 102, and concentricwith its longitudinal axis, is ultrasound probe 105. As probe tip 102compresses biological tissue, ultrasound probe 105 measures blood flowand tissue properties (soft tissue depth and density) in the localtissue region beneath the probe.

Turning now to FIG. 9, a single probe on a flexible arm is shown forcollecting anatomical, biomechanical and physiological data of atranstibial residual limb. FIG. 9 shows the single probe on a flexiblearm 96 conducting measurements on a residual limb 109 of a transtibialamputee patient. The following steps are taken to collect a full dataset of the residual limb.

Step 1.

In a first step, the skin strain and unloaded shape of the biologicalsegment is measured as a function of joint pose using the procedureoutlined previously. To this end, the biological limb is first markedwith a matrix of small (˜2 mm diameter), black-ink dots 110 across theentire skin-surface area for which the interface is designed tointeract. The specific anatomical location and distance between thesedots 110 need not be precise, but the resolution, or the number of dotsper cm² is important, as this resolution defines the resolution of theresulting skin strain field. Further, the pattern of dots should berandomized, providing the opportunity to create a unique skin specklepattern for each anatomical region. As an alternative to a matrix ofsmall dots, the skin of the biological segment of interest can bespeckled with a sponge where the sponge is first dipped into FDAapproved body paint. By dabbing the painted sponge across the skinsurface, a unique pattern of skin speckles can be created.

Step 2.

Next, separate poses, or joint postures of the biological segment ofinterest, are captured using the four cameras 101 on the single probe.For this step, the user of the probe grabs probe handle 103 andpositions the single probe with the cameras 101 pointed towards thebiological limb. With the limb held in a static position, the probe usertakes a total of ˜30 photographs from distinct probe positions, so as toimage all sides of the limb. During this exercise, probe tip 102 doesnot make contact with the biological segment. Using these ˜30 digitalphotographs for each limb pose, 3D models are generated for each limbposition. This exercise is then repeated for several limb poses. Forexample, for the case of the transtibial residual limb shown in FIG. 9,three knee positions could be images: 0, 45 and 90 degrees. As outlinedpreviously, these data are then used to generate both the unloaded shapeand skin strain of the biological segment as a function of anatomicallocation.

Step 3.

Next, the location-aware, force-sensitive probe 96 is used toestimate 1) the orthogonal impedance of body tissue at each anatomicalpoint (viscoelastic map), and 2) the stress or strain tissue thresholdwhere the subject first experiences discomfort at each anatomical point(sensitivity map). Single probe 96 is pushed against the subject's bodypart (e.g. residual limb 109) where the 3-D position of the probe tip102 is measured by the instrument at all times using high precisionencoders 100 and spring compression potentiometer 108. Additionally, ifthe body part under measurement is not stationary, the single probe 96also tracks its location in the same 3-D reference frame as themeasurement probe. The probe tip 102 is positioned at the surface of thehuman body in a perpendicular orientation to the surface area of theskin at each anatomical point, and the probe user(practitioner/clinician/user) then pushes with varying force,compressing the subject's tissues. Since the force-sensing probemeasures position, both of itself and of the biological part, both theviscoelastic and sensitivity maps are ascertained for the body part ofinterest.

The anatomical locations where tissue impedance and sensitivities aremeasured can be at each dot of the speckled pattern (e.g. 110 in FIG.9). In so doing, impedance/sensitivity measurements at all threevertices of each triangular element (e.g. in FIG. 5) can be determined.In so doing, tissue orthogonal stiffness, damping and stress/strainlevel where discomfort occurs can be correlated with the x, y, and z ofeach triangular node, where the peak skin strain is also known betweenall node points.

Since the pattern of skin speckles 110 is unique at each anatomicallocation, a single camera image from cameras 101 taken of a small regionof the skin surface can be used to determine the anatomical position atwhich the camera's lens is pointed or directed. Since the geometricposition of each camera is fixed relative to probe tip 102 andultrasound probe 105 with knowledge of probe spring 106 compression, theanatomical location of probe tip force application on the body can bedetermined. As noted previously, such an anatomical-positioningalgorithm is achieved by comparing the single, anatomically-local imageto the full speckle patterns across the entire biological segmentdetermined in Step 2. With such a positioning algorithm, the biologicallimb can move during the times when impedance measurements are not beingmade without having to measure such limb movements. However, during animpedance measurement the biological limb has to remain stationaryglobally, or if there is translational or rotational limb bone movement,such movements have to be measured, so as to determine accurately theamount of tissue compression caused by the probe force. Here thetranslation of the biological limb bone structure along the longitudinalaxis of the probe has to be subtracted from the measured probe tip 3-Dlocation upon tissue force application to determine an accurate measureof tissue compression and ultimately tissue impedance.

In addition, probe 96 uses ultrasound 105 to image the body to ascertaininternal tissue properties and blood flow and nervous tissuetransduction dynamics, and how such dynamics change as increasing forceis applied on the tissue by probe tip 102. Ultrasound transducer 105within probe tip 102 is used to gather very detailed tissue densitydata, soft tissue depth (orthogonal distance from the bodies surface tothe bone), and blood flow dynamics. Since single probe 96 measures forcesimultaneous with the ultrasound measurement, blood flow just beneaththe ultrasound probe 105 can be measured as a function of applied probeforce to determine how blood dynamics may be altered upon increasingapplied external pressure.

Finally, the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point is measured to produce asensitivity map. Here the subject verbally reports his/her level ofdiscomfort with each applied probe force for each anatomical point atwhich probe force is applied. When the subject first reports discomfortat each anatomical location, that applied force and tissue strain isrecorded and later used to create a sensitivity map.

Embodiment II Probe Array on a Flexible Arm

To increase the speed with which the body part of interest can beanatomically, biomechanically and physiologically mapped, multipleprobes can be employed. This embodiment represents a human-actuatedinstrument where the instrument's user, under their own muscle control,applies positions about, and forces onto, the biological segment ofinterest. In FIG. 10A, a probe array 115 is shown on a flexible arm 119.A close up of the probe array 115 is shown in FIG. 10B. In thisembodiment, a plurality of male probes 116 move linearly into and out ofa probe head 117. Inside probe head 117 are force sensors for measuringthe force applied on each male probe 125 when each male probe 125 ispushed against a biological segment, compressing tissue and applyingforce. In addition to force, each male probe 125 has an ultrasound probe126 mounted co-axially inside the tip of the male probe 125. Inaddition, a small camera 127 is mounted on the outside surface, ordiameter, of each male probe 125.

Flexible arm 119 comprises rotary joints (e.g. 121) so that each maleprobe 125 can be positioned anywhere in a large 3-D volume relative toflexible arm base 120. To accurately measure the 3-D position of the tipof each male probe 125, high precision encoders 122 contribute to theprecise measurement of the lab frame position of each male probe 125.

A practitioner/clinician/subject grabs handle 118 of the probe array andpushes the probe array against a body surface of interest where theprobes are approximately perpendicular to the body's surface. An exampleis shown in FIG. 11 where two probe arrays 130, 131 are pressed againstthe residual limb of a transtibial amputee 132. As shown, each probearray (right 130 and left hand 131) are co-axially aligned; for example,in the configuration shown in FIG. 11, the longitudinal axis of eachindividual probe 125 of the right hand probe array 130 is co-axiallyaligned with its corresponding individual probe within the left handprobe array 131. Through this co-axial alignment between the twoopposing probe arrays, the forces applied on the body segment sum tozero since the user of the probe arrays (clinician/practitioner/subject)exerts equal and opposite forces where all probe forces sum to zero.With the net external force applied equal to zero, the biological limb'seffective center of mass remains stationary, where the internal bones ofthe limb do not accelerate when the limb's external tissues are beingcompressed by the probe arrays. If the user of the probe array iscareful to balance applied forces from the array onto the biologicallimb, the limb will not translate and thus, measuring the position andorientation of the biological limb (e.g. 132) may not be necessary. Ifthere is a risk that the biological limb will translate and/or rotateduring tissue-compression experiments, technologies can be used tomeasure the limb's full orientation in 3-D space. For example, passivereflective markers can be positioned on the biological limb, and theindividual probe cameras 127 can measure the x, y, z locations of thesemarkers relative to lab frame in 3-D space. If during tissue compressionthe biological limb moves, this movement can be subtracted away from thex, y, z location of each probe's tip or point of tissue contact whenestimating the amount of tissue compression, and the speed of tissuecompression. Other technologies can be used to determine the position ofthe biological limb during probe measurements including, but not limitedto, electromagnetic signals, accelerometers, gyroscopes, magnetometers,active markers on the biological limb and receivers positioned off thelimb in lab frame, or any other location technology or combination oflocation technologies.

Still referring to FIG. 11, two probe arrays 130 and 131 are shown oneach on flexible arms used in the mapping of a residual limb 132 of atranstibial amputee patient. The following steps are taken to collect afull data set of a biological limb segment.

Step 1.

In a first step, the skin strain and unloaded shape of the biologicalsegment is measured as a function of joint pose using the procedureoutlined previously. To this end, the biological limb is first markedwith a matrix of small (˜2 mm diameter), black-ink dots across theentire skin-surface area for which the interface is designed tointeract. The specific anatomical location and distance between thesedots need not be precise, but the resolution, or the number of dots percm² is important, as this resolution defines the resolution of theresulting skin strain field. Further, the pattern of dots should berandomized, providing the opportunity to create a unique skin specklepattern for each anatomical region. As an alternative to a matrix ofsmall dots, the skin of the biological segment of interest can bespeckled with a sponge where the sponge is first dipped in FDA approvedbody paint. By dabbing the painted sponge across the skin surface, aunique pattern of skin speckles can be created.

Step 2.

Next, separate poses, or joint postures of the biological segment ofinterest, are captured using the cameras 127 on each probe. For thisstep, the user of the probe array grabs probe handle 118 and positionsthe probe array with cameras 127 pointed towards the biological limb.With the limb held in a static position, the probe array user takes atotal of ˜30 photographs from distinct probe positions, so as to imageall sides of the limb. During this exercise, the tip of probe 125 doesnot make contact with the biological segment. Using these ˜30 digitalphotographs for each limb pose, 3D models are generated for each limbposture. This exercise is then repeated for several limb poses. Forexample, for the case of the transtibial residual limb shown in FIG. 11,three knee positions could be images: 0, 45 and 90 degrees. As outlinedpreviously, these data are then used to generate both the unloaded shapeand skin strain of the biological segment as a function of anatomicallocation.

Step 3.

Next, the location-aware, force-sensitive probe array 115 is used toestimate 1) the orthogonal impedance of body tissue at each anatomicalpoint (viscoelastic map), and 2) the stress or strain tissue thresholdwhere the subject first experiences discomfort at each anatomical point(sensitivity map). Probe array 115 is pushed against the subject's bodypart (e.g. residual limb 132) where the 3-D position of the tip of eachmale probe 125 is measured by the instrument at all times using highprecision encoders 122, and the spring compression potentiometercorresponding to said male probe 125 housed within probe head 117.Additionally, if the body part under measurement is not stationary, theprobe array 115 also tracks its location in the same 3-D reference frameas the measurement probe array. The tip of each male probe 125 ispositioned at the surface of the human body in a perpendicularorientation to the surface area of the skin at each anatomical point,and the probe array user (practitioner/clinician/subject) then pusheswith varying force, compressing the subject's tissues. Since theforce-sensing probe array measures position, both of each individualmale probe 125 and of the biological part (e.g. 132), both theviscoelastic and sensitivity maps are ascertained for the body part ofinterest.

Since the pattern of skin speckles 110 (Step 1) is unique at eachanatomical location, a single camera image from each camera 127 taken ofa small region of the skin surface can be used to determine theanatomical position at which the camera's lens is pointed or directed.Since the geometric position of each camera is fixed relative to the tipof each male probe 125 and ultrasound probe 126 with knowledge of eachmale probe's relative compression distance within probe head 117, theanatomical location of the application of each male probe tip on thebody can be determined. As noted previously, such ananatomical-positioning algorithm is achieved by comparing the single,anatomically-local image to the full speckle patterns across the entirebiological segment determined in Step 2. With such a positioningalgorithm, the biological limb can move during the times when impedancemeasurements are not being made without having to measure such limbmovements. However, during an impedance measurement the biological limbhas to remain stationary globally, or if there is translational orrotational limb bone movement, such movements have to be measured, so asto determine accurately the amount of tissue compression caused by eachmale probe 125 force. Here the translation of the biological limb bonestructure along the longitudinal axis of each male probe 125 has to besubtracted from the measured male probe 125 displacement upon tissueforce application to determine an accurate measure of tissue impedance.

In addition, probe array 115 uses ultrasound 126 to image the body toascertain internal tissue properties and blood flow and nervous tissuetransduction dynamics, and how such dynamics change as increasing forceis applied on the tissue by the tip of each male probe 125. Ultrasoundtransducer 126 within each male probe tip 125 is used to gather verydetailed tissue density data, soft tissue depth (orthogonal distancefrom the bodies surface to the bone), and blood flow dynamics. Sinceprobe array 115 measures force applied to each male probe simultaneouswith the ultrasound measurement, blood flow just beneath each ultrasoundprobe 126 can be measured as a function of applied probe force todetermine how blood dynamics may be altered upon increasing appliedexternal force via each male probe 125.

Finally, the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point is measured to produce asensitivity map. Here the subject verbally reports his/her level ofdiscomfort with each applied probe force for each anatomical point atwhich probe force is applied. When the subject first reports discomfortat each anatomical location, that applied force and tissue strain isrecorded and later used to create a sensitivity map.

Although the probe array 115 can map a biological limb faster than thesingle probe of embodiment I, it has several disadvantages. First, sincethe probe array is planar where each probe tip is the same length atequilibrium when no force is applied, each probe would not apply a forcethat is perpendicular to the bodies' surface, especially when thebiological body's surface is highly curved. Second, a sensitivity map isdifficult to measure since reported discomfort by the subject cannot beprecisely attributed to an exact anatomical location or to a specificmale probe 125.

Embodiment III Finger Probe on a Flexible Arm

Embodiment III overcomes Embodiment II's lack of probe orthogonality andlocation specificity, while improving Embodiment I's speed with which abiological segment can be mapped anatomically, biomechanically andphysiologically. This embodiment represents a human-actuated instrumentwhere the instrument's user, under their own muscle control, appliespositions about, and forces onto, the biological segment of interest. InFIGS. 12A and 12B, a finger probe on a flexible arm is shown. Finger 135inserts into finger socket 136 that is connected to a flexible arm 137with high precision encoders 138 at each arm degree of freedom todetermine the global 3-D position in space of finger socket 136. Mountedon finger socket 136 is male probe 139 that moves linearly into, and outof, female probe housing 140. When male probe 139 compresses tissue, itpushes on a force sensor within the female probe housing 140.

The force sensor comprises a spring and linear potentiometer such asdescribed in Embodiment I, as best seen in FIG. 8B; spring 106 andpotentiometer 108). The probe measures force with a simple spring andlinear position measurement device. Through a measurement of thedeflection of the sensor's physical spring housed in the female probehousing 140, the force can be estimated using the force-deflectionrelation of the spring (e.g. F=−kx). It will be understood by those ofordinary skill in the art that the finger probe could also have a forcesensor that is either capacitive, strain-gauge based, or resistive. Inaddition, the finger probe also includes ultrasound to image the body toascertain internal tissue properties and blood flow and nervous tissuetransduction dynamics, and how such dynamics change as increasing forceis applied on the tissue. An ultrasonic transducer probe is co-axialwith male probe 139 and is used to gather very detailed tissue densitydata, soft tissue depth (orthogonal distance from the bodies surface tothe bone), and blood flow dynamics (e.g. how blood flow is altered uponincreasing applied external pressure). For photogrammetric data, a smallcamera 141 mounts on finger socket 136.

FIG. 13 shows finger probes on flexible arms collecting anatomical,biomechanical and physiological data of a transtibial residual limb. Twofinger probes, one for the left hand 142 and one for the right hand 143are preferably provided. Each probe is attached to a lab frame locationthrough a flexible arm base 144.

A known geometric relationship exists between the mounting location ofcamera 141 on finger socket 136 and the location of female probe housing140. Given the measured amount of insertion of male probe 139 intofemale housing 140, or probe force via the potentiometer recording, theprecise location of the tip of the male probe 139 to finger socket 136and camera 141 is known and can be recorded with a data acquisitionsystem. Further, via position sensing from precision encoders 138, thelocation of the finger socket 136 is known relative to flexible arm base144 of flexible arm 137 and can likewise be recorded with a dataacquisition system.

It should be understood by those of ordinary skill in the art thatcommunication wires from the finger probe force sensor, camera,ultrasound probe and encoders can travel through flexible arm 137 oradjacent the arm. Alternatively, an antenna can be positioned on thefinger socket 136 for wireless transmission of sensory data to areceiver within a data acquisition system comprising computer, A/Dconversion, signal conditioners, and power supply. Specifically, thephotogrammetric, force, position, velocity, acceleration and ultrasounddata from the finger probe can be communicated wirelessly or wired. Thewireless method can be IR-based, Bluetooth, or any other wirelesscommunication method such as an open electromagnetic frequency.

FIG. 13 shows two finger probes 142 and 143 each on flexible arms usedin the mapping of a residual limb 145 of a transtibial amputee patient.The following steps are taken to collect a full data set of a biologicallimb segment:

Step 1

In a first step, as with Embodiments I and II, the skin strain andunloaded shape of the biological segment is measured as a function ofjoint pose. To this end, the biological limb is first marked with amatrix of small (˜2 mm diameter), black-ink dots across the entireskin-surface area for which the interface is designed to interact. Thespecific anatomical location and distance between these dots need not beprecise, but the resolution, or the number of dots per cm² is important,as this resolution defines the resolution of the resulting skin strainfield. Further, the pattern of dots is randomized, providing a uniqueskin speckle pattern for each anatomical region. As an alternative to amatrix of small dots, the skin of the biological segment of interest canbe speckled with a sponge where the sponge is first dipped into FDAapproved body paint. By dabbing the painted sponge across the skinsurface, a unique pattern of skin speckles 110 can be created.

Step 2

Next, separate poses, or joint postures of the biological segment ofinterest, are captured using the camera 141 on each finger probe. Forthis step, the user of the finger probe positions the finger probe withcamera 141 pointed towards the biological limb. With the limb held in astatic position, the finger probe user takes a total of ˜30 photographsfrom distinct finger probe positions, so as to image all sides of thebiological limb. During this exercise, the tip of male probe 139 doesnot make contact with the biological segment. Using these digitalphotographs for each limb pose, 3D models are generated for each limbposture or pose. This exercise is then repeated for several limb poses.For example, for the case of the transtibial residual limb shown in FIG.13, three knee positions could be images: 0, 45 and 90 degrees. Asoutlined previously, these data are then used to generate both theunloaded shape and skin strain of the biological segment as a functionof anatomical location.

Step 3

Next, the location-aware, force-sensitive finger probe is used toestimate 1) the orthogonal impedance of body tissue at each anatomicalpoint (viscoelastic map), and 2) the stress or strain tissue thresholdwhere the subject first experiences discomfort at each anatomical point(sensitivity map). Finger probe is pushed against the subject's bodypart (e.g. residual limb 145) where the 3-D position of the tip of eachmale probe 139 is measured by the instrument at all times using both theforce sensor potentiometer and the high precision encoders 138.Additionally, if the body part under measurement is not stationary, thefinger probe also tracks its location in the same 3-D reference frame asthe flexible arm base 144. The tip of each male probe 139 is positionedat the surface of the human body in a perpendicular orientation to thesurface area of the skin at each anatomical point, and the probe arrayuser (practitioner/clinician/subject) then pushes with varying force,compressing the subject's tissues. Since the force-sensing finger probemeasures position, both of the tip of the male probe 139 and of thebiological part (e.g. 145), both the viscoelastic and sensitivity mapsare ascertained for the body part of interest.

Since the pattern of skin speckles (Step 1) is unique at each anatomicallocation across the biological segment, a single camera image frommounted camera 141 taken of a small region of the skin surface can beused to determine the anatomical position at which the camera's lens ispointed or directed. Since the geometric position of the camera 141 isfixed relative to the tip of male probe 139 (and correspondingultrasound probe) with knowledge of male probe's relative compressiondistance within the female probe housing 140, the anatomical location ofthe application of each male probe tip on the body can be determined. Asnoted previously, such an anatomical-positioning algorithm is achievedby comparing the single, anatomically-local image to the full specklepatterns across the entire biological segment determined in Step 2. Withsuch a positioning algorithm, the biological limb can move during thetimes when impedance measurements are not being made without having tomeasure such limb movements. However, during an impedance measurementthe biological limb has to remain stationary globally, or if there istranslational or rotational limb bone movement, such movements have tobe measured, so as to determine accurately the amount of tissuecompression caused by the male probe 139 force. Here the translation ofthe biological limb bone structure along the longitudinal axis of themale probe 139 has to be subtracted from the measured male probe 139global, lab frame displacement upon tissue force application todetermine an accurate measure of tissue impedance.

In addition, the finger probe of Embodiment III uses ultrasound to imagethe body to ascertain internal tissue properties and blood flow andnervous tissue transduction dynamics, and how such dynamics change asincreasing force is applied on the tissue by the tip of each male probe139. The ultrasound transducer within the tip of male probe 139 is usedto gather very detailed tissue density data, soft tissue depth(orthogonal distance from the bodies surface to the bone), and bloodflow dynamics. Since the finger probe measures force applied to the maleprobe 139 simultaneous with the ultrasound measurement, blood flow justbeneath each ultrasound probe, along the projection of longitudinal axisof male probe 139 into the biological segment, can be measured as afunction of applied finger probe force to determine how blood dynamicsmay be altered upon increasing applied external force.

Finally, the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point is measured to produce asensitivity map. Here the subject verbally reports his/her level ofdiscomfort with each applied probe force for each anatomical point atwhich probe force is applied. When the subject first reports discomfortat each anatomical location, that applied force and tissue strain isrecorded and later used to create a sensitivity map.

In increase the speed with which a biological segment can be mapped,many finger probes can be employed up to 10 finger probes, one for eachfinger on the right and left hands of the finger probe user. A pluralityof finger probes can map a biological limb faster than the single probeof embodiment I, without the disadvantages of Embodiment II, namely poororthogonality and specificity. Since each finger probe is controlled bya biologically-actuated finger, orthogonality can be achieved where eachfinger probe applies a force perpendicular the body's surface. Inaddition, since each biological finger is independent in its forceapplication and position, the stress/strain threshold where a subjectexperiences discomfort at a specific anatomical location can bedetermined.

Other advantages of the finger probe include its spatial versatility andproximity to the biological segment. Since the finger probe has sensorslocated directly on each finger tip, it has great spatial versatility;difficult areas of the body can be mapped where there is little spacefor a bulky instrument. For example, if the upper thigh needs to bemapped in the case of a transfemoral prosthetic socket or legexoskeleton, the finger probe can readily take measurements in themedial crotch area. The finger probe also has an improved proximity tothe biological member compared to other types of instruments such asEmbodiment I and II. Since the force, ultrasound, and photogrammetricsensors are located on the biological finger tip, the distance from thefingers to the biological segment being mapped is relatively small,allowing the user of the instrument to more readily palpate thebiological member during data collection.

Embodiment IV Untethered Finger Probe

The spatial versatility of Embodiment III is improved over Embodiments Iand II, but the design is still not optimal. Because each finger socketis tethered to a flexible arm in Embodiment III, versatility may belimited since the flexible arm may cause obstructions when mapping somedifficult-to-reach body segments. Further, the flexible arm makes theinstruments of Embodiment I, II, and III somewhat bulky and difficult totransport.

As a resolution to these difficulties, an untethered finger probe isshown in FIGS. 14A and 14B. This embodiment represents a hybrid-actuatedinstrument comprising both synthetic actuation and human-poweredactuation. The probe 150 comprises a hemispherical external finger cap152 and a hemispherical internal finger cap 159. The external andinternal caps, 152 and 159, respectively, are hemispherical at theirterminus end, or tip, because: 1) a hemispherical tissue indenter can bereadily modeled using finite element modeling software, and 2) when thebody is touched lightly by the external cap 152, it interfaces the bodywith a small surface area or point, making the determination of thecontact point more tractable.

The external 152 and internal 159 finger caps are separated by adielectric elastomer device, or force sensor, 160. While dielectricelastomers have often been used for actuation and power generation, theycan also be used as an integrated force sensor. When a dielectricelastomer device (dielectric elastomer material, such as silicone, withimbedded compliant electrodes) is mechanically deformed, both thecapacitance and dielectric resistance of the material is changed. Thus,compliant electrodes will be embedded within the dielectric elastomerdevice 160 to measure mechanical forces applied to the external fingercap 152 when a subject's tissues are being compressed. Relative to theX, Y, and Z coordinate frame of the external finger cap 152, a forcevector can be measured having force components in the X, Y, and Zdirections. Here Z is perpendicular to the external finger cap outwardlydirected along the longitudinal axis of the untethered finger probe 150,and X and Y are orthogonal to this longitudinal Z axis. Specifically, asan example, when a force is applied to external finger cap 152 having ageneral direction along the longitudinal Z axis of the finger probe 150,the dielectric elastomer device 160 compresses, becoming thinner andundergoing a capacitance change that correlates with an applied Z force.Alternatively, the untethered finger probe 150 could exert a shear forceagainst a biological segment, resulting in a shear force applied to theexternal finger cap 152, or a force in the XY plane. Such a shear forcewould cause the dielectric elastomer device 160 to compress with adistinct strain field compared to the strain field caused by the pureZ-axis force of the previous example. Electrode patterning upon thedielectric material of device 160 using microfabrication is designed todifferentiate between forces applied in X, Y and Z directions. Such aforce sensor 160 offers several potential advantages over traditionalsensors including operation over large strain ranges, ease of patterningfor distinctive sensing capabilities, flexibility to allow uniqueintegration into components, stable performance over a wide temperaturerange, and low power consumption.

The dielectric elastomer device 160 comprises silicone positionedbetween patterned electrodes, one patterning near the inner surface ofthe layer 160 (in close proximity to the internal finger cap 159), and asecond patterning on its outer surface (in close proximity to theexternal finger cap 152). Such a dielectric sensor measures changes incapacitance when the silicone material is compressed under an externallyapplied pressure, or force, applied to external finger cap 152. Withinthe walls of layer 160, conductive traces pass from each electrode toprocessing unit 155 via external finger cap 152 (not shown). Withoutloss of generality, finger probe 150 could also have a force sensor thatis resistive, piezoelectric based, strain-gauge based,spring-potentiometer based, or any other force sensing technology.

The untethered finger probe 150 also includes ultrasound to image thebody to ascertain internal tissue properties, blood flow and nervoustissue transduction dynamics, and how such dynamics change as increasingforce is applied on the tissue by external finger cap 152. An ultrasoundtransducer probe 153 is used to gather very detailed tissue densitydata, soft tissue depth (orthogonal distance from the bodies surface tothe bone), and blood flow dynamics (e.g. how blood flow is altered uponincreasing applied external tissue pressure). Ultrasound probe 153 ismechanically grounded to hemispherical external finger cap 152. When aforce is exerted on the external finger cap 152, ultrasound probe 153moves through a clearance hole 158 within the internal finger cap 159.Within the walls of the external finger cap 152, wires pass from theultrasound probe 153 to processing unit 155 (wiring not shown).

The untethered finger probe 150 also comprises a small camera 154 forthe collection of photogrammetric data. Camera wires pass from camera154 to processing unit 155 (wiring not shown).

Untethered finger probe 150 also comprises a full inertial measurementunit (IMU) 161. The IMU is attached to external finger cap 152, andcomprises 3 accelerometers, 3 rate gyros and a magnetometer.

Processing unit 155 includes, but is not limited to, a microprocessor,RAM, A/D conversion, USB port, and power supply. Data from the forcesensor 160, ultrasound probe 153, camera 154, and IMU 161 aretransmitted to processing unit 155 where basic signal conditioning isperformed such as A/D conversion, filtering, amplification, etc. priorto wireless transmission. In one embodiment, during data collectionprocessed sensory data are wirelessly transmitted via antenna 156 to areceiver located on a data acquisition station not attached to thefinger probe, but in the vicinity of the probe. It will be understood bythose of ordinary skill in the art that the photogrammetric, force,position, velocity, acceleration and ultrasound data from the fingerprobe can be communicated wirelessly using IR-based, Bluetooth, or anyother wireless communication method such as an open electromagneticfrequency. In another embodiment, sensory data are stored on processingunit 155, and later transferred via the USB port to a computer forprocessing and modeling. In this framework, processing unit 155 hassubstantial memory for data storage, so that wireless or wired transferof data could be completed when it is convenient to do so. In apreferred embodiment, memory storage is sufficient to store all the datafor at least a complete map of an entire body segment anatomically,biomechanically and physiologically. With such storage space, thetransfer of data can occur subsequent to the data collection on thehuman subject, enabling a greater degree of convenience. This type offramework is critical for data collections that occur in very remoteregions of the world, where the transport of a computer is inconvenient.In this type of situation, the untethered finger probe, or probes, couldbe carried in a backpack to any remote village in poor areas of theworld. Subjects could be mapped, and the data set could later beuploaded from the finger probe, or probes, for analysis and design of amechanical interface.

Finally, untethered finger probe 150 comprises an actuator controlledand powered by processing unit 155. A vibration actuator 151, such as apager motor, is mechanically grounded to the external finger cap 152.When the vibration actuator is activated, its motor spins an asymmetricmass that causes the finger probe 150 to vibrate. When finger probe 150is in contact with a biological segment, these vibrations cause a forceripple against the tissue directly beneath the external cap's tissueforce application (center of pressure point on external cap 152 due totissue contact). The dielectric elastomer device 160 and processing unit155 records such force ripples, and the IMU 161 and processing unit 155records the accelerations caused by the vibrations. The accelerationdata are then low pass filtered by processing unit 155 to estimate themaximum tissue compression and time rate of change of compression.During tissue palpation, the measured force signal from force sensor 160is combined with the tissue compression, and rate of compression data,to estimate viscoelastic tissue properties directly at the point ofapplication, or center of pressure, between finger probe 150 and thesubject's tissue.

To increase the speed with which a biological segment can be mapped,multiple untethered finger probes can be employed. FIGS. 15A and 15Bshow different views of five untethered finger probes on each digit ofthe hand in the form of a data acquisition glove 170. Each finger probecomprises all the elements shown in FIG. 14 including vibratory actuator151, external finger cap 152, ultrasound probe 153, camera 154,processing unit 155, antenna 156 and IMU 161. It will be understood bythose of ordinary skill in the art that a single, central processor andantenna, located on the back of the hand or palm, could be used insteadof individual processor units 155 and antenna's 156 on each finger tip.For the case where the computation and wireless transmission hardwarewere located on the back of the hand, electrical transmission wireswould run down the fingers from each individual finger probe 150 to thecentral processor and antenna.

FIGS. 16A and 16B show different view of ten untethered finger probesworn on the left and right hand in the form of two gloves, one for theleft hand 170 and a second for the right hand 171. As an example, theten untethered finger probes in FIG. 16 are used to rapidly map abiological ankle-foot complex 172 that has been speckled 173 with FDAapproved body paint. The following steps are taken to collect a fulldata set of any biological limb segment, including for example theankle-foot complex shown in FIGS. 16A and 16B.

Step 1

In a first step, as with Embodiments I, II and III, the skin strain andunloaded shape of the biological segment is measured as a function ofjoint pose. To this end, the biological limb is first marked with amatrix of small (˜2 mm diameter), black-ink dots across the entireskin-surface area for which the interface is designed to interact. Thespecific anatomical location and distance between these dots need not beprecise, but the resolution, or the number of dots per cm² is important,as this resolution defines the resolution of the resulting skin strainfield. Further, the pattern of dots is randomized, providing a uniqueskin speckle pattern for each anatomical region. As an alternative to amatrix of small dots, the skin of the biological segment of interest canbe speckled with a sponge where the sponge is first dipped into FDAapproved body paint. By dabbing the painted sponge across the skinsurface, a unique pattern of skin speckles can be created.

Step 2

Next, separate poses, or joint postures of the biological segment ofinterest, are captured using the camera 154 on each untethered fingerprobe. For this step, the user of the untethered finger probes positionthe finger probes with cameras 154 pointed generally towards thebiological limb. With the limb held in a static position, the fingerprobes' user, or probe operator, takes photographs from distinct fingerprobe positions, so as to image all sides of the biological segment (˜30or more photographs for each limb pose). To improve image quality, alight flash can be used with each camera 154, and/or a continuous lightoutput source from each camera, so as to minimize problematic shadowsthat complicate subsequent data processing. During this exercise, theuntethered finger probes do not make contact with the biological segment(e.g. Left image in FIG. 16). Using these digital photographs for eachlimb pose, 3D models are generated for each limb posture or pose. Thisexercise is then repeated for several limb poses. For example, for thecase of the ankle-foot complex shown in FIG. 16, three ankle positionscould be imaged: 0 degrees, dorsiflexion 15 degrees and plantar flexion20 degrees, as well as three subtalar joint positions: 0 degrees,inversion 10 degrees, and eversion 10 degrees. As outlined previously,these data are then used to generate both the unloaded shape and skinstrain of the biological segment for each segment pose as a function ofanatomical location.

Using the untethered finger probe 150, an alternate method for measuringthe unloaded shape of a biological segment is through the use of the IMU161 and force sensor 160 on each finger probe 150. In this method, theoperator of the finger probe 150 starts from a single point marked onthe skin in the region of the biological segment of interest. The fingerprobe operator then moves his finger gently along the surface of thebiological segment, with the hemispherical external cap 152 lightlytouching the skin surface. During this movement, a position trajectoryis computed along those skin points contacted by the external finger cap152. Specifically, in this method the IMU 161 is used first to estimatethe lab frame spatial trajectory (X, Y, Z positions versus time) of theIMU 161 located on the finger probe 150 by performing a zero velocityupdate when the finger probe is held stationary (zero accelerationexcept for gravity) at the starting point marked on the skin, and thenintegrating forward. By integrating forward, the lab frame X, Y, Z IMUtrajectory in 3-D space relative to the starting point is computed.

After this IMU trajectory calculation is performed, the estimate of thelab frame X, Y, Z trajectory of the external finger cap contact pointagainst the skin, or center of pressure, is computed by conducting ageometric transformation from the lab frame IMU X, Y, Z trajectory tothe measured center of pressure location on the external finger cap 152.In this calculation, the center of pressure position relative to thefixed position of the IMU 161 on the finger probe 150 is computed usingthe force sensor 160 and the fixed position of the IMU relative to theexternal cap 152. This local frame position trajectory of the center ofpressure point relative to the IMU is then added to the lab frametrajectory of the IMU 161, to compute the lab frame trajectory of thecenter of pressure point as the finger probe 150 is moved across theskin surface. By repeating this finger movement pattern at a highresolution of skin points, always starting from the same starting point,the shape of the limb can ultimately be determined. To minimizeintegration drift error from the IMU calculation, the operator's fingermovement along the skin surface is done quickly at high velocitystarting from a zero-velocity starting point on the skin surface.

It will be understood by those of ordinary skill in the art that manyfinger movement patterns could be employed to map the shape of abiological limb. For example, the operator could first move his fingerto key anatomical points, and then subsequently map the shape of theskin surface relative to these anatomical locations. If for example,there were N anatomical locations geometrically distributed across thebiological segment, the operator could map N skin surface regionsimmediately adjacent each anatomical location. These N surface regionscould later be stitched together computationally to form the overallshape of the biological segment.

To improve upon the speed with which the operator of the untetheredfinger probe 150 can measure a biological segment's unloaded shape,multiple finger probes in the form of a data acquisition glove 170, canbe used (See FIGS. 15A and 15B). For example, when all ten digits of theoperator employ finger probes 150, the operator would place each of hisfingers with a finger probe 150 at a marked starting point on the skinsurface of a biological segment of interest, forming hand grasp postureson the biological segment. From that starting posture, the operatorwould then move his fingers gently across the surface of the skin,following the biological segment's contours. Using the same IMU 161 andforce sensor 160 calculation, the X, Y, Z lab frame position trajectoryof each finger probes' center of pressure would be computed, designatingthe contour of the biological segment beneath each finger probe'spathway across the skin. By repeatedly returning to the same left andright hand grasp postures, and repeating this finger movement pattern ata high resolution of skin points, the shape of the limb can ultimatelybe determined. To spatially couple each finger probe's skin trajectoryto its neighboring finger probe trajectories within a single left orright hand glove, the position in 3D space of each finger probe 150relative to adjacent finger probes needs to be measured. For thismeasurement, the IMU 161 on each finger probe 150 is required to measurethe orientation (pitch, roll, yaw) of each finger probe 150. Further,additional sensors are required on the data acquisition glove 170. Forexample, additional IMU's placed on the middle or proximal phalanx ofeach digit are necessary to measure the hand grasp posture, and thelinear distance between adjacent worn finger probes 150. Alternatively,by using dielectric material to form the data acquisition glove 170,wherein electrode traces are fabricated onto the dielectric material ofthe glove using a microfabrication methodology to form capacitivestretch sensors passing around the finger digits, measurements of digitflexion/extension and abduction/adduction could be made, and suchmeasurements could be used to compute the linear distances betweenadjacent finger probes.

Step 3

Next, each untethered finger probe is used to estimate 1) the orthogonalimpedance of body tissue at each anatomical point (viscoelastic map),and 2) the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point (sensitivity map). Eachuntethered finger probe is first pushed against the subject's body part(e.g. ankle-foot complex 172 in right image of FIG. 16) where thelongitudinal axis of the finger probe is approximately perpendicular tothe surface of the body at the point of contact. Using the untetheredprobe, or probes, tissue impedance can be estimated in one of two ways.

In a first method, the vibration actuator 151, is used to cause eachfinger probe 150 to vibrate. When the vibration actuator is activated,its motor spins an asymmetric mass that causes the finger probe 150 tovibrate. When finger probe 150 is in contact with a biological segment,these vibrations cause a force ripple against the tissue directlybeneath the external cap's tissue force application (center of pressurepoint on external cap 152 due to tissue contact). The dielectricelastomer device 160 and processing unit 155 records such force ripples,and the IMU 161 and processing unit 155 records the accelerations causedby the vibrations. The acceleration data are then low pass filtered byprocessing unit 155 to estimate the maximum tissue compression and timerate of change of compression. Subsequent to tissue palpation, themeasured force signal from force sensor 160 is combined with the tissuecompression, and rate of compression, to estimate tissue mechanicalimpedance for that tissue region underlying the hemispherical externalcap 152 of each respective finger probe or probes. This estimationcomputation is performed using finite element modeling to capture thecontinuous viscoelastic nature of biological tissue.

In a second method, the IMU 161 is used to estimate a change in tissuestate (position and speed) by performing a zero velocity update when thefinger probe is held stationary (zero acceleration except for gravity)against the tissue and then integrating. An LED on processing unit 155is used to inform the probe user of the zero velocity update status.Once the zero velocity update is complete, the LED turns green from red,and the user of the finger probe then quickly pushes against the tissue,applying a greater force and tissue compression. By integrating forwardthe change in probe position in 3-D space can be estimated, and if thebiological segment does not translate, the amount of tissue compressioncan be determined. Simultaneous to this estimate of tissue compression,and compression rate, the finger probe measures the applied force on thebiological segment. Subsequent to tissue palpation, the measured forcesignal from force sensor 160 is combined with the tissue compression,and rate of compression, estimate from the IMU 161 calculation toestimate tissue mechanical impedance for that tissue region underlyingthe hemispherical external cap 152 of each respective finger probe orprobes. This estimation computation is performed using finite elementmodeling to capture the continuous viscoelastic nature of biologicaltissue. It is important to estimate tissue impedance using tissuecompression data from only the first moments after the zero velocityupdate. Preferably only the first 0.3 seconds of data after the zerovelocity update should be used, as later times would result in tissuecompression errors that are too great due to drift in the displacementestimate.

Since the pattern of skin speckles (Step 1) is unique at each anatomicallocation across the biological segment, a single camera image frommounted camera 154 taken of a small region of the skin surface can beused to determine the anatomical position at which the camera's lens ispointed or directed. Since the geometric position of camera 154 is fixedrelative to the external finger cap 152 (and corresponding ultrasoundprobe 153), the anatomical location of the application of each fingerprobe on the body can be determined (assuming the probe user pushes onthe body in a direction that is perpendicular to the body's surface). Asnoted previously, such an anatomical-positioning algorithm is achievedby comparing the single, anatomically-local image to the full specklepatterns across the entire biological segment determined in Step 2. Withsuch a positioning algorithm, the biological limb can move during thetimes when impedance measurements are not being made without having tomeasure such limb movements. However, during an impedance measurementthe biological limb has to remain stationary in a global sense if thebiological segment's global position is not being tracked or measured.

A practitioner/clinician/subject using the instrument pushes each fingerprobe against a body surface of interest where each probe's orientationin contact with the body is approximately perpendicular to the body'ssurface at the point of probe force application. An example is shown inFIG. 16 where two 5-digit finger probe systems, or gloves 170 and 171are shown. Here several finger probes are pressed against the subject'sankle-foot complex 172. As shown, the right hand 171 has a finger probethat exerts a force equal to but opposite the finger probes of the lefthand 170. Through this apposing alignment between the two probe gloves,the forces applied on the body segment sum to zero since the user of thegloves, or probe operator, exerts equal and opposite forces where allfinger probe forces sum to zero. With the net external force appliedequal to zero, the biological limb's effective center of mass remainsstationary, where the internal bones of the limb do not accelerate whenthe limb's external tissues are being compressed by the finger probes.If the user of the finger probes is careful to balance applied forcesfrom the probes onto the biological limb, the limb will not translateand thus, measuring the position and orientation of the biological limb(e.g. 172 in FIGS. 16A and 16B) may not be necessary. If there is a riskthat the biological limb will translate and/or rotate duringtissue-compression experiments, technologies can be used to measure thelimb's full orientation in 3-D space. For example, passive reflectivemarkers can be positioned on the biological limb, and a camera orcameras can be mounted on the finger probe glove (170, 171 in FIGS. 16Aand 16B) in order to measure the x, y, z locations of these passivemarkers relative to lab frame in 3-D space. If there istranslational/rotational limb bone movement during impedance/sensitivitymeasurements, such movements have to be measured, so as to determineaccurately the amount of tissue compression caused by the finger probeforce. In this case, the translation of the biological limb bonestructure along the longitudinal axis of the finger probe 150 has to besubtracted from the measured finger probe 150 displacement upon tissueforce application to determine an accurate measure of tissue impedanceand sensitivity strain threshold. Other technologies can be used todetermine the position of the biological limb during probe measurementsincluding, but not limited to, electromagnetic signals, accelerometers,gyroscopes, magnetometers, active markers on the biological limb andreceivers positioned off the limb in lab frame, or any other locationtechnology or combination of location technologies.

In addition, the finger probe of Embodiment IV uses ultrasound to imagethe body to ascertain internal tissue properties and blood flow andnervous tissue transduction dynamics, and how such dynamics change asincreasing force is applied on the tissue by finger probe 150. Theultrasound transducer 153 is used to gather very detailed tissue densitydata, soft tissue depth (orthogonal distance from the bodies surface tothe bone), and blood flow dynamics. Since the finger probe measuresforce applied to the external finger cap 152 simultaneous with theultrasound measurement, more accurate ultrasound data can be acquired.Since the ultrasound signal changes with applied force, or pressure,between the ultrasound head and the tissue being imaged, ultrasound datacan be compared between distinct anatomical points at a fixed level ofapplied force, increasing the consistency and repeatability of theultrasound data. This combination of force sensing and ultrasoundsensing also enables the probe operator to measure blood flow justbeneath each ultrasound probe, along the projection of longitudinal axisof finger probe 150 into the biological segment, as a function ofapplied finger probe force to determine how blood dynamics may bealtered upon increasing applied external force.

Finally, the stress or strain tissue threshold where the subject firstexperiences discomfort at each anatomical point is measured to produce asensitivity map. Here the subject verbally reports his/her level ofdiscomfort with each applied probe force for each anatomical point atwhich probe force is applied. When the subject first reports discomfortat each anatomical location, that applied force and tissue strain isrecorded and later used to create a sensitivity map.

As shown in FIGS. 16A and 16B, to increase the speed with which abiological segment can be mapped, many finger probes can be employed,for example, up to 10 finger probes, one for each finger on the rightand left hands of the finger probe user. A plurality of finger probescan map a biological limb faster than the single probe of embodiment I,without the disadvantages of Embodiment II, namely poor orthogonalityand specificity. Since each finger probe is controlled by abiologically-actuated finger, orthogonality can be achieved where eachfinger probe applies a force perpendicular the body's surface. Inaddition, since each biological finger is independent in its forceapplication and position, at each anatomical location the stress/strainthreshold when a subject first experiences discomfort can be determined.

Other advantages of the untethered finger probe include its spatialversatility, its proximity to the biological member being mapped, andits ease of transport. Since the finger probe has sensors locateddirectly on each finger tip, and given the fact that the probe isuntethered, affords it great spatial versatility; difficult areas of thebody can be mapped where there is little space for a bulky, tetheredinstrument with a flexible arm. For example, if the upper thigh needs tobe mapped in the case of a transfemoral prosthetic socket or legexoskeleton, the untethered finger probe can readily take measurementsin the medial crotch area without risk that the flexible arm willinterfere, or block data collection in some way. The finger probe alsohas an improved proximity to the biological member compared to othertypes of instruments, e.g. Embodiment I and II. Since the force,ultrasound, IMU and photogrammetric sensors are located on thebiological finger tip, the distance from the fingers to the biologicalsegment being mapped is relatively small, allowing the user of theinstrument to more readily palpate the biological member during datacollection. Finally, the untethered finger probe is readilytransportable; without the need for a flexible arm or tether, the fingerprobe, or probes, could be thrown into a backpack, for example, andemployed to map the biological segment of a subject located in a remotearea of the world.

Step 2: Mapping Biological-Limb Model Representation to MechanicalInterface Shape and Viscoelastic Properties

Mapping Skin-Strain Model to the Tensile Viscoelastic Properties of theMechanical Interface

Understanding how the skin is stretched as a body segment is moved isparamount to mechanical interface design. As an example, in the case ofa transtibial leg amputation, FIGS. 5A, 5B and 6A, 6B, 6C clearly showrelatively large longitudinal skin strain at, and just proximal to, thepatella, as well as large circumferential strains proximal to the kneejoint when the knee assumes a flexed posture. Using conventionalprosthetic socket technology, an amputee typically wears a liner that isrolled across the residual limb. By making the coefficient of staticfriction high between the skin and liner materials, designers haveeffectively lowered relative movement at that interface, reducinguncomfortable rubbing and chaffing. However, current liner technologydoes not comprise continuously varying tensile material properties thatare informed by a skin-strain model as described in the previoussection. Consequently, in areas of large skin strain, inflexibility inthe liner causes skin discomfort due to high skin shear stresses imposedby the liner material. For example, in the case of a transtibialamputation, inflexibility in the liner in the high strain regions, orthe patella and proximal knee areas, cause skin discomfort, especiallywhen an amputee sits with knees flexed for an extended period of time.

In one embodiment of the present invention, we propose a liner thatapplies minimal shear stress on the skin when the biological segmentchanges posture, minimizing discomfort at the skin-interface junction.To achieve this goal, the mechanical strain energy stored within theliner is minimized when the biological limb is moved to a pose withlarge skin strains. We achieve this goal by continuously adjusting thetensile viscoelastic properties of the material spatially across theliner surface.

As an example, for the case of a transtibial amputation as shown in theskin-strain model of FIGS. 6A, 6B, 6C, large tensile skin strains areclearly visible longitudinally at, and proximal to, the knee patella. Inthis region of the residual limb, the skin-strain triangles arestretched longitudinally, or along the long axis of the thigh,indicative of the skin being under a large tensile stretch in thatdirection (detail B in FIG. 6C). In this region, the liner should bemore stiff along the directions of minimum strain, indicated by the bluevectors, and less stiff along the red vectors representing maximumstrain. This would serve to support the knee around the patella butpermit knee flexion. In addition, due to muscle contractions upon kneeflexion, large tensile skin strains are clearly visiblecircumferentially in the region of the leg proximal to the knee joint(detail A in FIG. 6B). Here, the proposed liner should permitcircumferential expansion of the limb and be stiffer along the thigh'saxial direction. The corresponding liner material adjacent to theselarge skin strain directions would be fabricated with aproportionally-small stiffness and damping, so as to minimize the amountof shear forces against the skin when the knee is flexed. In thisinvention, we teach of using a quantitative mapping from the skin-strainmodel to the corresponding tensile viscoelastic properties of theadjacent liner.

In the skin-strain model described in the previous section, a lineconnects each black-dot to an adjacent black-dot. In the modelingmethodology, a strain is computed for each of these dot-to-dot lines,forming a whole grid of interconnected triangles (See FIG. 6A). In oneembodiment of the present invention, the stiffness of the adjacent linermaterial to tensile stretch is numerically computed along the linebetween each set of two black-dot points, or each leg of a skin-straintriangle. The numerical relationship could be linear or nonlineardepending upon the type of mechanical interface, the region of the bodyfor which an interface is to be constructed, and the specific needs ofthe user. In one embodiment, the mapping from the skin-strain model tothe liner tensile viscoelastic properties is linear; liner stiffnessalong each leg of a skin-strain triangle is inversely proportional tothe computed maximal skin strain, namely, where the skin strain islarge, the corresponding tensile liner stiffness is small. Further,where the skin strain is small, the corresponding tensile linerstiffness is large. In one embodiment, in regions of large skin strain,a black-dot to black-dot stiffness equal to zero could be preferable, oralternatively a small stiffness that does not cause skin discomfort whenthe joint is held at a large-strain pose for an extended period of time.It will be understood by those of ordinary skill in the art that suchproposed relationships between skin strain and adjacent syntheticmaterials of a mechanical interface can be applied to any wearablegarment, shoe or device. For example, an athletic shoe of this inventionwould comprise an inner sock liner with continually varying modulusproperties directly correlated to the underlying skin strain valuesassociated with joint movements.

Mapping the Biological-Limb Shape-and-Impedance Model to MechanicalInterface Shape-and-Impedance Properties: A Linear Model

The human anatomy is complex and consists of multiple materials ofdifferent properties. For example, a transtibial residual limb consistsof bones, (femur, tibia, fibula, and the patella), muscles (tibialis,gastrocnemius, peroneus longus, etc.) and other anatomical landmarksincluding, but not limited to, the tibial tuberosity, medial femoralcondyle, lateral femoral condyle and the medial tibial flare.

In one embodiment of the present invention we employ a quantitativemapping between the viscoelastic properties of the body when the body iscompressed orthogonal to the skin surface, and the correspondingproperties of the mechanical interface.

For areas on the body for which an interface is to be designed, theunderlying anatomical components and their viscoelastic properties arequantitatively related to the stiffness and damping of the adjacentmechanical interface. For one embodiment of the present invention, wewill have an interfacing material adjacent to each anatomical locationwith inverse stiffness and damping characteristics to that of the body.Although an inverse linear mapping algorithm is used here, there couldexist a nonlinear mapping including but not limited to parabolic,hyperbolic, trigonometric, exponential functions, and differentialequations will create unique spatial material compositions within themechanical interface for each anatomical location. The available toolsare limited to automatically measure the body's stiffness and dampingproperties when a residual limb is compressed perpendicular to its skinsurface. As such, in one embodiment of the present invention, we assumethat the gross stiffness and damping properties of the body scale to thesoft tissue depth at that anatomical point. Here soft tissue depth isdefined as the orthogonal distance between the surface of the skin andthe intersection of bone tissue when the body is not being compressedand is in a state of equilibrium. For boney protuberances such as thefibula head in the transtibial residual limb, the soft tissue depth issmall and the body is stiff to orthogonal compression. In distinction,in the calf region the soft tissue depth is relatively larger and thebody is relatively softer to orthogonal compression.

In one embodiment, the perpendicular distance from the skin surface tothe bone obtained from MRI or other imaging data is used as a grossestimate of the body's viscoelastic properties. FIG. 17 shows thequantitative relationship between mechanical interface stiffness, ordurometer, and body stiffness represented as the percentage of softtissue depth. Here the horizontal axis is the soft tissue depth, D,normalized by the maximal soft tissue depth, D_(max), multiplied by 100.Both linear and non-linear curves are presented showing the possiblevariation in the relationship between interface durometer andcorresponding soft tissue depth. Generally, as soft tissue depthdecreases, and body stiffness increases, the adjacent interface becomesincreasingly soft. Where there are boney protuberances, the adjacentinterface will be soft and compliant, but where the body is soft with alarge soft tissue depth, the adjacent interface is designed to be morerigid.

Another critical parameter that describes the mechanical interfacedesign is the percent of soft tissue compression, namely the percentchange in the soft tissue depth caused by the interface during anon-loaded state. In FIG. 18, the percent of soft tissue compression isplotted vertically, and the percent of tissue depth is plottedhorizontally. Here the horizontal axis is the soft tissue depth, D,normalized by the maximal soft tissue depth, D_(max), multiplied by 100.Further, the vertical axis is the soft tissue compression caused by theinterface, normalized by the maximum soft tissue compression, multipliedby 100. Several linear curves are shown, depicting that as soft tissuedepth increases, the amount that the interface compresses the tissueincreases. Although only linear curves are shown in FIG. 18, additionalembodiments could include nonlinear relationships such as parabolic,hyperbolic, trigonometric, exponential functions, and differentialequations. Generally, where the body is soft, or where soft tissue depthis high, the interface will compress the tissues more. Where there is aboney protuberance, and the body is stiff with a small soft tissuedepth, the interface will compress the tissues by a small amount or notat all. Such an inverse relationship between body stiffness and tissuecompression results in a more uniform pressure field across the residuallimb surface.

It will be understood by those of ordinary skill in the art that thelevel of tissue compression by the mechanical interface may depend uponanatomical location. For example, when there are underlying nerves andvessels that may be more sensitive to external pressure, the level oftissue compression by the interface will have to be reduced accordingly.A single curve mapping the level of tissue compression to bodyviscoelastic properties may not be universally applied across the entirebiological segment, but may vary as a function of anatomical location.Clearly, a plurality of curves may be required to fully capture thequantitative mapping between tissue compression levels, bodyviscoelastic properties and anatomical location.

Mapping the Biological-Limb Shape-and-Impedance Model to MechanicalInterface Shape-and-Impedance Properties: An Optimization Procedure

In the previous embodiment, linear mappings (FIGS. 17 and 18) were oftenassumed, relating the output of the shape-and-impedance biomechanicalmodel to a numerical description of the interface's shape and impedanceproperties. In this section, a mathematical optimization framework ispresented for defining the mapping that does not assume linearity apriori. The procedure employs the digital anatomical data of that partof the body for which an interface design is sought, to attain thatinterface shape and impedance that produces a uniform interface pressureapplied to the biological limb, and a minimized spatial pressuredifferential in the presence of atrophy by the biological limb.

Before presenting the optimization procedure, we define key variables:

-   -   A. From a set of digital points {right arrow over (S)}_(i)        ^(v)(X,Y,Z) located on the surface of the biological limb to be        interfaced with a mechanical device, create a 3D volume. Here Z        is in the direction of the gravitational vector, whereas X and Y        are perpendicular to the Z-axis and to each other.    -   B. From three neighboring points or vertices {right arrow over        (S)}₁ ^(v)(X,Y,Z), {right arrow over (S)}₂ ^(v)(X,Y,Z), and        {right arrow over (S)}₃ ^(v)(X,Y,Z), define the area vector        ({right arrow over (A)}_(i)) of each triangle, within the grid,        directed outwardly and orthogonally from the surface of the        biological limb. Note the origin of area vector {right arrow        over (A)}_(i)(X,Y,Z) is located at the center of area at point        {right arrow over (S)}_(i)(X,Y,Z).    -   C. Define the unit area vector as {right arrow over        (e)}_(i)={right arrow over (A)}_(i)/A_(i), or the area vector        divided by the magnitude of the area vector. This unit vector is        directed outwardly and orthogonally from the center of area of        the section defined by the three neighboring vertices {right        arrow over (S)}₁ ^(v)(X,Y,Z), {right arrow over (S)}₂        ^(v)(X,Y,Z), and {right arrow over (S)}₃ ^(v)(X,Y,Z).    -   D. Define the angle θ_(i) between the line of the unit area        vector and the vertical Z-axis.    -   E. Define the total area at the top of the socket in the Z        direction, or A_(Z_top). A simplified approach to estimate        A_(Z_top) is to assume a circle defining a plane that is        orthogonal to direction Z, with a diameter equal to the average        diameter of the residual limb adjacent the socket's upper, or        most proximal, brim or cutline. More rigorously, A_(Z_top) is        the total area in the Z direction of the adjoining surface        connecting the line around the residual limb surface at the        upper, or most proximal, brim or socket cutline.    -   F. Calculate the uniform Pressure (P_(uni)) within the        prosthetic socket. It is approximated as P_(uni)=W/A_(Z_top) for        a transtibial or transfemoral socket for a person in quiet,        single-leg standing with body weight W. Alternatively, as a        worst case, one could assume a uniform pressure equal to        3W/A_(Z_top). Here the factor of 3 is an estimate of the dynamic        loading experienced during running.    -   G. Calculate the vector force (F _(i)) parallel but oppositely        directed from area vector ({right arrow over (A)}_(i)) from the        uniform socket pressure (P_(uni))        {right arrow over (F)} _(i) =−P _(uni) *{right arrow over (A)}        _(i)    -   H. Determine the residual limb impedance I_(i) with stiffness        K_(i) and damping B_(i) components of each node point {right        arrow over (S)}_(i)(X,Y,Z) at the center of area {right arrow        over (A)}_(i) (impedance is based on the mechanical properties        of skin, muscle, fat and bone measured in the direction of the        applied Force vector, {right arrow over (F)}_(i)=−P_(uni)*{right        arrow over (A)}_(i))    -   I. Calculate {right arrow over (r)}_(i)(ΔX, ΔY, ΔZ) to get the        new point {right arrow over (S)}_(i)(X,Y,Z)*. The 3D volume from        the set of points {right arrow over (S)}_(i)(X,Y,Z)* determines        an optimal shape of the socket at load {right arrow over        (F)}_(i)=−P_(uni)*A _(i) that achieves a uniform-socket,        residual-limb interface pressure.        -   1. {right arrow over (r)}_(i)={right arrow over            (S)}_(i)(X,Y,Z)*−{right arrow over (S)}_(i)(X,Y,Z)        -   2. For one embodiment, we estimate {right arrow over            (r)}_(i)(ΔX, ΔY, ΔZ) by assuming a linear approximation for            body stiffness, or K_(i)=C_(i)*d_(i) where d_(i) is the            scalar soft tissue depth defined as the distance from the            center of area at {right arrow over (A)}_(i) on the surface            of the residual limb at point {right arrow over            (S)}_(i)(X,Y,Z) to the surface of the bone measurable using            MRI, and C_(i) is a proportionality constant between body            stiffness K_(i) and the distance d_(i). Thus, {right arrow            over (r)}_(i)={right arrow over (F)}_(i)/(C_(i)*d_(i)).

Optimization

The procedure thus far estimates the shape of the residual limb {rightarrow over (S)}_(i)(X,Y,Z)* under a uniform pressure, P_(uni), with aload at each node equal to {right arrow over (F)}_(i)=−P_(uni)*A _(i)and the amount of tissue compression at that load, orY_(i)=−P_(uni)*{right arrow over (A)}_(i)/(K₁). Using a simplified modelfor estimating body stiffness K_(i)=C_(i)*d_(i), we have {right arrowover (r)}_(i)=−P_(uni)*{right arrow over (A)}_(i)/(C_(i)*d_(i)). SinceP_(uni) W/A_(Z_top), {right arrow over(r)}_(i)=−(W/(A_(Z_top)C_(i)d_(i)))*({right arrow over (A)}_(i)).However, what is still unknown is the optimal interface impedance, orfor a static load assuming quiet standing, the optimal interfacestiffness k_(i). In this example, the damping force term b_(i)*{rightarrow over (V)}_(i) is not a consideration since it is a statics problemwith tissue compression velocity {right arrow over (V)}_(i) equal tozero. To optimize the stiffness of the socket interface k_(i) at eachinterfacing node {right arrow over (S)}_(i)(X,Y,Z)* at pressure P_(uni)that yields a constant socket pressure in a variable-impedance socket,we minimize the pressure differential (δP/δZ), or the change ininterface pressure along the surface of the residual limb in the Zdirection in the presence of an atrophy or hypertrophy disturbance.

-   -   A. The socket interface stiffness k_(i) describes the stiffness        of the interface adjacent to node i.    -   B. The amount of interface elastic compression at node i is        equal to:        {right arrow over (s)} _(i) ={right arrow over (F)} _(i) /k        _(i)=(−P _(uni) *{right arrow over (A)} _(i))/k _(i)=(−W/A        _(Z_top) *{right arrow over (A)} _(i))/k _(i)    -   C. Consider that the residual limb has changed shape at the zero        load condition from {right arrow over (S)}_(i)(X,Y,Z) to S_(i)        ^(d)(X,Y,Z) due to residual limb atrophy or hypertrophy. We can        define an atrophy or hypertrophy disturbance vector {right arrow        over (a)}_(i) as        {right arrow over (a)} _(i) ={right arrow over (S)} _(i)        ^(d)(X,Y,Z)−{right arrow over (S)} _(i)(X,Y,Z).    -   D. In one embodiment, the disturbance vector is equal to:        {right arrow over (a)} _(i) =−D _(i) *d _(i) *{right arrow over        (e)} _(i)    -    where {right arrow over (e)}_(i)={right arrow over        (A)}_(i)/A_(i) defined earlier, d_(i) is the soft tissue depth        defined earlier, and D_(i) is a proportionality constant. We        assume here that the atrophy or hypertrophy disturbance is        orthogonal to the residual limb surface at node i, and is        proportional to the soft tissue depth at that point.    -   E. After the disturbance, the interface spring compression would        be:        -   1. {right arrow over (T)}_(i)={right arrow over            (s)}_(i)−{right arrow over (a)}_(i)−ΔZ_(i)({right arrow over            (g)}/g) and the force at node i would be {right arrow over            (F)}_(i)=k_(i)[{right arrow over (s)}_(i)−{right arrow over            (a)}_(i)−ΔZ_(i)({right arrow over (g)}/g)]        -   2. Here ΔZ_(i)=[W−Σ_(i)[k_(i)({right arrow over            (s)}_(i)−{right arrow over (a)}_(i))·{right arrow over            (g)}/g]]/[Σ_(i)[k_(i) cos θ_(i)]]        -   3. After the disturbance, the pressure field is no longer            uniform, and is equal to:            P _(i) ={right arrow over (F)} _(i) /{right arrow over (A)}            _(i)        -   4. Minimize the pressure differential

$\frac{\partial P_{i}}{\partial Z}$

-   -   -    in the Z direction along the surface of the body from node            to adjacent node by varying node stiffnesses k_(i)        -   5. For the array of interface stiffnesses k_(i) ^(min) that            minimize

$\frac{\partial P_{i}}{\partial Z^{\min}},$

-   -   -    identify S_(i)(X,Y,Z)** that gives the new interface            equilibrium (unloaded) shape, or            -   a. S_(i)(X,Y,Z)**={right arrow over (s)}_(i)+{right                arrow over (S)}_(i)(X,Y,Z) where            -   b. {right arrow over (s)}_(i)={right arrow over                (F)}_(i)/k_(i)=(−P_(uni)*{right arrow over                (A)}_(i))/k_(i) ^(min)=(−W/A_(Z_top)*{right arrow over                (A)}_(i))/k_(i) ^(min)

Step 3: Mechanical Interface Fabrication

The most advanced prototyping and CAM technology on the market will beused to seamlessly integrate spatially-varying viscoelastic propertiesinto the mechanical interface design. It is understood by those ofordinary skill in the art that the final mechanical interface can bemanufactured using both traditional and state-of-the-art methodsincluding, but not limited to, casting, 3D printing, mechanical linkagesof disparate materials and shape deposition manufacturing.

Fabrication of Tensile Viscoelastic Properties

It will be understood by those of ordinary skill in the art that linerviscoelastic properties can be varied spatially in a number of ways,including but not limited to, varying liner thickness, density, materialcomposition and type, and/or material structure (e.g. through the use ofsmall material hinges across the liner surface).

In one embodiment, liner thickness is varied to accomplish spatialviscoelastic variation. Here each strain triangle leg (as an example seeFIGS. 6A-C) has a corresponding thickness of the liner inverselyproportional to the maximum skin-strain computed. In another embodiment,the numerical mapping computes the average of the three skin strainscorresponding to each leg of a skin-strain triangle (an example is shownin FIGS. 5A, 5B), and then an inversely-proportional relationshipdefines the corresponding liner thickness at that triangular region.

In another embodiment, a plurality of different material types areemployed within the liner. Along each leg of a skin-strain triangle forwhich large strains occur, a thin compliant material is employed withinthe liner, while adjacent the small-strain leg of a skin-strain trianglea separate material is attached to further increase the liner thicknessand stiffness in such regions. For example, in the transtibial residuallimb case, shown in FIGS. 6A-C, for the area proximal to the knee jointthe skin is stretched circumferentially but not longitudinally along thelong axis of the thigh upon knee flexion. The adjacent liner couldcomprise of a thin compliant material spanning the entire region, andattached to it strips of added material running longitudinal to the longaxis of the thigh. When the thigh muscles contract and expand upon kneeflexion, and the skin stretches circumferentially, the thin, compliantliner material would accommodate this stretch with minimal shear forceapplied to the skin, while the longitudinal strips would add structuralintegrity to the liner interface. In distinction, for the patella, andthe region just proximal to the patella, shown in FIG. 6, the skinstretches longitudinally but not circumferentially as the knee assumes aflexed posture. In such regions, the thin strips of added material wouldrun circumferentially, while the underlying thin, compliant materialwould connect adjacent strips, allowing the skin to stretchlongitudinally upon knee flexion with minimal shear stress applied tothe skin.

Fabrication of Compression Viscoelastic Properties

Various methods have been suggested to relieve pressure over bonyprotuberances and other anatomical landmarks in passive prostheticsockets. In conventional approaches, different materials have beenbonded or mechanically attached together to relieve pressure onanatomical protrusions. Other CAD/CAM methodologies include the use ofdouble walls, and most recently, the creation of mechanical compliantfeatures in a 3-D printing process.

In one embodiment of the present invention we employ variable impedancesseamlessly integrated into socket production using advanced 3D printingtechnology. 3D printing has been used in design of medical technologiesfor decades. However, the methodologies and capabilities of the machineshave continued to evolve. Objet Geometries Inc. (North America, 5Fortune Drive, Billerica, Mass. 01821, USA, T: +1-877-489-944) producesthe most advanced 3D printer that uses their PolyJet Matrix™ Technology.In FIG. 19, the Objet Connex500 is shown. This technology enablesdifferent material durometers to be simultaneously jetted in theproduction of the same mechanical interface, allowing for spatiallyvarying viscoelastic properties across the interface surface. Witha16-micron, high-resolution print layer, high dots-per-inch in both Xand Y resolution, and an easy-to-remove support material property, thistechnology is ideal for the development of prosthetic and orthoticprototypes.

There is a relatively large library of standard materials used by theConnex family of 3D printers. In addition, composite materials can becreated to produce Digital Materials™ to give a wide range of materialproperties; a desirable feature in prosthetic and orthotic designsmapped from calculated biological limb stiffness and damping properties.

Shown in FIGS. 20A-20P is an example of how an Objet 3-D printingprocess can be employed in the fabrication of a prosthetic socketprototype for a transtibial amputee. In the first (FIGS. 20A-20D) andsecond row (FIGS. 20E-20H), MRI images and corresponding soft tissuedepth models are shown for the right leg of a transtibial amputee.Orientation from left to right for all images are anterior, lateral,medial and posterior, respectively. Acquired MRI data are used to designthe varying viscoelastic features within the socket wall.

As with FIGS. 7A-C, the second row FIGS. 20E-20H shows different viewsof the soft tissue depth model of the residual limb. As defined earlier,the soft tissue depth is the orthogonal distance D between the skinsurface and a bone intersection. Here, red regions show large tissuedepths, yellow regions moderate depths, and green regions relativelysmall depths. For these depth models, the patella tendon was removed,exposing the soft tissue depth in the region of the patella tendon justdistal to the patella (shown as the red region in the left-most image).

In the third row in FIGS. 20I-20L, different views of a 3-D printedprosthetic socket is shown where every material color corresponds to amaterial having a distinct durometer and tensile strength. Here, the redmaterial has the highest durometer and tensile strength, while the greenmaterial has the smallest durometer and tensile strength. Morespecifically, Table 1 in FIG. 31 shows the mapping from soft tissuedepth to interface material tensile strength. All these distinctcompression viscoelastic features are integrated together seamlessly sothat the sockets are manufactured in one piece with limited postprocessing requirements. The color mapping is used above with softtissue depth being shown in millimeters (mm), and socket tensilestrength in mega-pascals (MPa).

In the fourth row of FIG. 20.4, the socket's most rigid, high tensilestrength material (shown in red in the third row) is modeled using anFEA analysis to evaluate structural integrity for vertical loadscomparable to that which would be experienced during standing andwalking. FIG. 21 shows the Von Mises Stress distribution andcorresponding color code used in FIG. 20.4. Assuming a 3X body weightvertical loading, the wall thickness of the red material shown in thethird row was varied to achieve an acceptable level of material stress.Additionally, the two struts, or bars, that connect the patella tendonregion of the socket to the distal socket base are included to achievestructural integrity; without these struts, the socket would be underrisk of collapsing upon vertical loading when the amputee stood orwalked with the socket interface.

Referring now to FIG. 22, the linear relationship used in the socketdesign and fabrication of FIGS. 20A-20P is shown. Here the quantitativemapping of interface modulus (plotted vertically) to soft tissue depth(plotted horizontally) is plotted, showing numerically how the interfacebecomes softer and softer as the body becomes stiffer and stiffer (withsmaller and smaller soft tissue depths). More specifically, FIG. 22shows the mapping between the Young's Modulus of socket interfacematerials shown in the third row of FIGS. 20I-20L to the soft tissuedepth at each location shown in the second row of FIGS. 20E-H, where itis color coded by categories of soft tissue depth.

Manufacturing for Durability

The fabrication example shown in FIGS. 20A-20P is problematic becausethe Objet 3-D printed material is unstable, degrading in time withunfavorable mechanical properties. In this section, we propose afabrication method that result in a more stable interface product.

From the optimized set of material impedances (k_(i)), atransformational mapping is established for manufacturing usingconventional processes including, but not limited to, molding, casting,shape deposition, and carbon composite lamination. In FIG. 23, atranstibial socket 50 is shown where each color represents a distinctmaterial durometer or impedance. Such a variable-impedance socket layercan be fabricated using shape deposition processes or by modulatingsilicone durometer spatially using standard silicone fabricationprocedures. The outer transparent element 52 is designed to transferload from the variable-impedance socket distally, while still allowingdeformation of the compliant regions of the socket 50. This outerelement 52 can be made of carbon fiber and is used to ensure structuralintegrity while allowing flexibility in the regions where compliance isneeded.

The ideal stiffness set k_(i) for the mechanical interface can beproduced with a spatially-varying impedance socket and integrated liner,encased in an outer carbon composite exoskeletal frame. In oneembodiment of the present invention, a liner 54, or a thin polyurethaneor silicone skin-tight sock, is bonded directly to the multi-materialsocket (See FIGS. 23 through 26), or can be attached and removed easilyin a donning and doffing process using standard attachment means such asa mechanical pin lock. In another embodiment of the invention, the linerand socket shown in FIGS. 24 through 26 could be fabricated as a singlepiece using polyurethane in a shape deposition process, or urethaneusing standard urethane fabrication strategies, or silicone.Alternatively, the liner 54 and socket 50 could be fabricatedseparately, and then attached subsequent to fabrication. Still further,in another embodiment the inner surface of the variable-impedance lineradheres to the body's skin using a synthetic “gecko” material thatincreases the shear strength between the skin and the interface, whilestill allowing easy donning and doffing of the artificial interface.

More specifically, in FIG. 24, a thin compliant material, or liner 54,is bonded at its distal aspect to the multi-material prosthetic socket50 shown in FIG. 23 to form a fully-integrated mechanical interface withthe body. On its inner surface 56, the liner 54 adheres intimately tothe human body using a synthetic gecko material. The liner system,because of its intimacy with the body, and its continuity with thesocket, fundamentally solves the issue of suspension within sockets,eliminating the need for a pin suspension or the like.

In FIG. 25, a full view of the multi-material prosthetic socket with theinternal liner 54 and socket 50 bonded within and a carbon fiber outermaterial 52.

In FIG. 26, the internal liner 54 bonded to the inner surface 58 of thevariable-impedance socket 50 within a carbon fiber element 52. Adjacentthe compliant regions within the socket 50, the carbon fiber element 52is spaced from the outer surface of the socket 50, so as to allowexpansion of the socket 50 in those regions.

In FIG. 27, the outer carbon fiber element 52 is designed for structuralintegrity. The external element could be made from carbon fiber or anystructural material capable of transferring loads to the externalprosthetic limb. The gap size between the variable-impedance socket andthe carbon fiber outer element is inversely proportional to the socket'sstiffness, namely, where the socket 50 is relatively soft, the gap islarge, allowing the compliant region of the socket 50 to expand outwardupon load bearing. In distinction, where the socket 50 is relativelystiff, the gap is made relatively small, so as to allow the transfer ofload from the variable-impedance socket to the carbon fiber externalframe element.

It will be understood by those of ordinary skill in the art that thevariable-impedance liner 54 and socket 50 in FIG. 26 could be fabricatedby using a single material having a spatially-variable geometry, or byusing multiple-material types. For example, the liner 54 of FIG. 26could be fabricated with a uniform thickness but with multiple durometersilicones. Further, the socket 50 of FIG. 26 could be fabricated with acontinuously-varying silicone thickness; where compliance is sought thesilicone would be relatively thick, whereas where stiffness is sought,the silicone would be relatively thin. Such a silicone socket 50 wallthat achieves a continuously-variable perpendicular modulus (correlatedto the perpendicular tissue modulus at each anatomical point) bycontinuously varying it silicone wall thickness has many advantages,including a simple manufacturing process with stable and durablematerials.

Fabrication of a Liner with Embedded Sensors

Continuous monitoring of physiological information within the socketliner can quantitatively inform socket fabrication and modification forthe improvement of socket fit and comfort. In addition, such technologywill provide previously unprecedented levels of information about thewearer's intent to aid in external bionic limb control. To achieve thislevel of monitoring, sensing electronics should be integrated into theliner itself. Wirelessly relaying sensed information from the linermaterial adjacent the residual skin to the external prosthetic socketelement is ideal in order to avoid needing wires and electricalconnectors passing from liner to external socket. There are two keyrelated problems for integrating sensing capability into the compliantliner. First, there is the packaging problem of actually placing thefront-end sensor (e.g. electromyographic (EMG) electrodes, force,pressure, shape, ultrasound, temperature, etc.) into the correctlocation relative to the body without causing discomfort orinconvenience to the wearer. Because such sensing modalities ideallyshould be located at the socket-skin interface, an ideal solution forpackaging would be the integration of compliant, miniature, and wirelesselectronics near the location of the sensor-to-body interface, formingan inner liner membrane that is smooth, continuous, and with skin-likemechanical properties.

The second problem is power consumption in the electronics. Clearly, theelements of the sensors that are in the liner or liner equivalent mustbe very small. If the electronics are integrated alongside the sensingelements, then the small size limits the amount of available energystorage or available harvested power. A target size for integrateddevices of a few cm² on a flexible substrate with minimal thicknesswould limit the energy budget to the range of 10's to 100's of Joules,2-3 orders of magnitude less than a cell phone battery. The main problemis that off the shelf electronics for physiological monitoring,processing, and wireless communication tend to consume many milliWattsof power (10s of mWs for most commercial radios). This power problem isprohibitive to integrating wireless electronics in the socket linerwithout a dramatic reduction in circuit power consumption.

In this section, the design and fabrication strategy of a liner (linerlayer 54 shown in FIG. 26) is described that employs a dielectricelastomer. While dielectric elastomers have often been used foractuation and power generation, they can also be used as an integratedsensor. When a dielectric elastomer device (dielectric elastomermaterial with imbedded compliant electrodes) is mechanically deformed,both the capacitance and dielectric resistance of the material ischanged. Thus, compliant electrodes are microfabricated within thedielectric elastomer liner material to measure mechanical forces (normaland shear) applied to the residual limb, and residual-limb, volume andshape changes. In addition, compliant electromyographic (EMG) sensorsare embedded within the liner for sensing muscle activity for thecontrol of active bionic joints.

In FIG. 28, a line drawing of the liner 1 is shown depicting thelocation of sensors and the electronic board 2 where sensing modalitiesare embedded within liner 54. The board 2 houses the liner electronicsand includes EMG amplification, A/D conversion, signal conditioning, andan antenna for wireless communications. Embedded within the wall of thesilicone liner are all sensors, electronics, and wiring such that theinner surface of the liner is smooth and continuous so as not to causediscomfort to the wearer. EMG sensors are soft, flexible electrodesshown at 5, 6, 11, 12, and 13. A commercial electrode can be used, e.g.from SmartTrace (1015663—SmartTrace Electrodes, Hospital Version).Electrodes are positioned over the muscles of the residual limb. Musclelocations can be determined from residual limb MRI or other imagingdata. EMG electrode 4 is positioned on the patella 3 for grounding.Before donning the liner, the residual limb is sprayed with abio-compatible adhesive for improved adhesion between the electrodes andskin.

Stretch sensing dielectric elastomers are used within the liner for themeasurement of forces applied on the residual limb from the liner andsocket. In addition, sensors are positioned within the liner to measurecircumferential shape changes. When such a device, (elastic polymer withcompliant electrodes), is mechanically deformed, both the capacitance ofthe device, as well as the electrode and dielectric resistance, arechanged. Such a sensor offers several potential advantages overtraditional sensors including operation over large strain ranges, easeof patterning for distinctive sensing capabilities, flexibility to allowunique integration into components, stable performance over a widetemperature range and low power consumption.

Components 7, 8, and 14 are representative patterned electrodes for themeasurement of liner forces in both normal and shear directions. Herethe silicone for force sensing is positioned between two electrodes, onebeneath the material layer, and a second on top of the material lay,forming a sandwich where the electrodes are the “bread pieces” and thesilicone is the “meat”. Such a dielectric sensor measures changes incapacitance when the silicone material is compressed under an externallyapplied pressure, and/or stretched causing the distance between theelectrodes to become smaller. Additionally, a dielectric sensor is usedto measure circumferential changes of the residual limb. Bands 9 and 10are stretch dielectric sensors using the dielectric sensor approachdescribed earlier. Within the walls of the silicone liner, wires ormicrofabricated conductive traces pass from each electrode to processingboard 2 (wiring not shown in FIG. 28). Sensory data are then transmittedto a receiver located on the external prosthesis.

EMG sensing is employed as the basis for controlling and modulating theresponse of a powered prosthesis. The liner is designed based upon suchEMG control requirements. The EMG electrodes are placed on the liner ina way to detect dorsiflexor (eg. Tibialis Anterior) and plantar flexor(eg. Soleus and Gastrocnemius) muscle activity in the residualtranstibial limb with these being used to signal the movement intent ofthe wearer. For instance, the EMG activity may signal intent to positionthe ankle into a dorsiflexed or plantar flexed position (jointequilibrium), to point the prosthetic foot upwardly while ascendingstairs or a hill, or to point the prosthetic foot downwardly so as tolift the wearer within reach of an object, or to point the toe whilewalking down stairs, for instance.

The EMG activity detected in the dorsiflexors may also signal the needfor increased stiffness and damping, together the impedance, in lateswing or early stance as might be needed to absorb energy (brake) whilewalking down a steep hill, for instance. The EMG activity detected inthe plantar flexors during mid-to-late stance may signal intent to walkfast, to run or to walk cautiously down a steep hill or stairs. Here,rather than controlling the ankle angle directly, the EMG activity canbe used to modulate the gain of the positive-torque feedback reflexresponse in the ankle prosthesis in accordance with wearer intent. Toaccomplish the modulation functions noted above, we transmit to theexternal active prosthesis the EMG signal from all five locations at therate of 125 Hz.

As noted earlier, the liner employs dielectric capacitance-based sensorsas the basis for quantitative measurement of socket force and limbshape/volume over time. Without loss of generality, such sensors canalso be used to monitor heart rate that will inform the level ofexertion during the day and over time. This information is logged in thenon-volatile memory of the prosthesis so that the clinician can observethe historical record of force (stress) and residual limbexpansion/contraction (residual limb shape/circumference) to inform theneed for an intervention—a socket modification or a new socket, forinstance.

As described in the previous subsection, the liner described hereincomprises dielectric capacitance-based sensors to measure limbcircumference and force. These transducers comprise force sensors andstretch sensors each using a capacitance measurement method to inferaxial or longitudinal deflection. The EAP capacitance, C_(EAP), isdefined as follows:

$C_{EAP} = \frac{{\epsilon\epsilon}_{0}A}{d}$

where ∈_(o) is the dielectric in a vacuum and c is the dielectricconstant of the polymer liner. A is the sensor electrode cross-sectionalarea defined as the product of the longitudinal dimensions, l_(x) andl_(y), and d is the thickness of the dielectric between the electrodesas shown in FIG. 29, which is an illustration of and electro-activepolymer electrode.

For the force sensing application, the capacitance change arises fromthe compression of the dielectric, decreasing the thickness, d. For theliner application, carbon black electrodes are printed roughly 10 mm×10mm with a sensor thickness of approximately 100 μm thick. Assuming adielectric constant of about 3, the nominal (zero deflection)capacitance will be roughly 30 pF. For the stretch application (upper)an electrode will be printed approximately 150 mm long and 10 mmwide—yielding a capacitance of approximately 130 pF. The correspondinglower stretch band is approximately half the length of the upper band,the capacitance will be roughly 65 pF. Assuming a typical deflection of10% in both the force and stretch applications, sensing circuitry isrequired that can detect capacitance changes with a precision of betterthan 1%, or about 0.25 pF.

To accomplish this, circuitry is employed to measure changes in the timeconstant, τ_(EAP), using an AC drive (˜20 kHz square wave for thepressure sense and ˜4 kHz for the stretch sense) with large sourceimpedance (˜1 MΩ). The 0.25 pF precision is achieved through detectionof the time constant change, δτ_(EAP), of approximately 250 nsec, orabout 1/200 of the excitation period. Two methods are proposed formaking the capacitance measurement. One employs an RMS measurement ofthe differential signal, using the fact that the RMS value of the signalshould be proportional to

$\frac{1}{\tau_{EAP}}.$In the other method, a comparator and low-pass filter is applied tomeasure implicitly the pulse-width of the comparator output that isproportional to τ_(EAP). In either case, the duty cycle of themeasurement is quite low since the measurement would be made 5-10 timesper day when the wearer is standing but not moving. Here an RMSmeasurement of socket pressure at all twenty locations and both upperand lower stretch could be made over a period of a second and thenreported to the external prosthesis where it could be stored innon-volatile memory. Such would give the clinician a daily historicalrecord from which trends relating to goodness-of-fit could be discerned.Further, force and circumference sensory information, in addition to EMGsensing, could be used by the controller of the external bionic limb forthe detection of gait phase, speed, terrain variation, and volitionaluser intent.

Although a prosthetic liner was described herein, it will be understoodby those of ordinary skill in the art that any apparel, shoe,prosthesis, orthosis, or exoskeleton could employ these inventive steps.For example, an athletic shoe that comprises these same design featureswould comprise a liner, or sock, made from skin-like dielectric materialwith embedded force, shape, EMG, temperature and ultrasound sensing. Thefoot liner would have a spatially varying tensile modulus correlated tothe underlying skin strain values caused by ankle flexion/extension andsubtalar inversion/eversion. The foot liner would support electronicsfor signal conditioning, A/D conversion, and wireless communication to areceiving station on the outer layer of the shoe, a wristband, anelectronic smart phone, or device. The variable-impedance intermediateshoe layer (corresponding to the socket 50 in FIG. 26) comprisesspatially-varying viscoelastic properties correlated to the viscoelasticproperties of the underlying ankle-foot tissues for perpendicular tissuecompressions. Finally, the external, or outer shoe layer (correspondingto the carbon fiber element 52 in FIG. 26) would comprise an elasticcomposite material for the storage of elastic strain energy during footstrike in walking and running.

Moreover, FIG. 28 shows a liner design with integrated force,circumference and EMG sensors. Liner manufacture will comprise twoseparate measurement and fabrication steps: 1) measure the biologicalsegment unloaded shape and the skin strain field at distinct jointposes; 2) design and fabricate a variable-compliant silicone linerhaving a thickness at each anatomical point that is inversely related topeak skin strain with embedded sensors and accompanying wire leads orconductive traces.

Step 1

Using standard photogrammetric tools, a model of skin strain as afunction of anatomical location and joint pose is generated. Such amodel is necessary to understand how the mechanical interface shouldmove and stretch relative to the skin surface, so as to minimize shearforces and discomfort at the skin-interface junction. In this procedure,the biological limb is first marked with a matrix of small (˜2 mmdiameter), black-ink dots across the entire skin-surface area for whichthe interface is designed to interact. The specific anatomical locationand distance between these dots need not be precise, but the resolution,or the number of dots per cm² is important, as this resolution definesthe resolution of the resulting skin strain field. In addition, theresolution can be variable, providing the opportunity to furtherinvestigate deformation in certain areas. Next, separate poses, or jointpostures of the biological segment of interest, are captured usingphotogrammetric tools. Using approximately 30 digital photographs foreach limb pose, 3D models are generated. The coordinates of the blackdots on the skin will then be marked and exported for analysis. Thepoint clouds for each pose will be triangulated in a correspondingmanner so the mapping of points to triangles is the same.

The black dots are the nodes of a finite element model and serve as thevertices for a surface triangulation. The deformation of each triangularelement from one pose to another are then decomposed into a translation,rotation, and stretch via an affine transform. FIGS. 30A and 30B showsan example where the equivalent strain of each triangulation resultingfrom the deformation of the original, extended pose to two differentlevels of knee flexion, respectively. The average strain is a scalarvalue that is useful for assessing the overall stretch of a skinelement. The average strain of each triangular face is analyzed andmapped to a color. Skin strain levels are shown for the partially flexedpose (FIG. 30A) and the fully flexed pose (FIG. 30B). Here higheraverage strain is shown around the knee patella due to the right pose'sincreased knee flexion.

Step 2

After the subject's skin strain has been measured, a variable-compliantsilicone liner is fabricated having a tensile modulus at each anatomicalpoint that is inversely proportional to the measured peak skin strain.Specifically, along directions where there is large skin strain, theadjacent liner will have a relatively low tensile modulus, whereas alongskin directions where the skin strain is small the adjacent liner willhave a relatively high tensile modulus. By varying the tensilecompliance of the liner in this manner, shear forces are minimized atthe liner-skin interface to mitigate skin damage and discomfort.

To fabricate such a variable-compliant liner, with integrated force,circumference and EMG sensors and their accompanying wire leads, a moldis 3-D printed having a negative space where silicone material is pouredand allowed to cure. A male plug is 3-D printed with a shapecorresponding to the unloaded biological segment of interest minus ˜4 mmcircumference reduction to achieve an appropriate liner tissuecompression once fabricated. Further, as part of the same 3-D printingprocess, a female mold is fabricated around the male plug such that thegap separating the female and male 3-D printed parts is equal to theliner thickness. After printing, the wire leads, or conductive traces,and sensor volumes are placed into the mold prior to liner fabrication.For example, the EMG sensors are attached on the outer surface of themale plug at regions of residual limb musculature where EMG signal canbe readily measured. Additionally, the grounding EMG electrode (e.g. EMGsensor 4 in FIG. 28) is placed on a boney protuberance of the male plug.Once the silicone material is poured within the gap separating male andfemale parts and has had adequate time to cure, the final silicone linerwith embedded sensors and lead wires is removed from the mold andcleaned for use. To vary tensile modulus across the liner spatially, theliner thickness at each point could be varied, or the durometer of thesilicone, or both. For the fabrication of a variable-durometer liner,separated cavities, or pockets, could be 3-D printed between the maleand female molds, and within each cavity, silicone, with a distinctdurometer, would then be injected and given time to cure.

In view of the above, an instrument for determining the anatomical,biomechanical, and physiological properties of a body segment thatincludes one or more force sensitive probes is provided. A humanoperator actuates one or more force sensitive probes, wherein the forcesensitive probes are positioned at the surface of the body segment. Theoperator pushes on the force sensitive probes with varying force appliedon the body segment to measure tissue deflection forces. The instrumentmay include one or more of gyroscopes, accelerometers, and magnetometerscapable of measuring changes in tissue deflection caused by the forcesensitive probes relative to a grounded reference frame in 3-D space,wherein the tissue deflection force data and the change in tissuedeflection data are used to compute segment tissue viscoelasticproperties. The instrument may be untethered or wireless.

It would be appreciated by those skilled in the art that various changesand modifications can be made to the illustrated embodiments withoutdeparting from the spirit of the present invention. All suchmodifications and changes are intended to be covered by the appendedclaims.

What is claimed is:
 1. A system for determining anatomical,biomechanical, and physiological properties of a body segment, includingan instrument, the instrument comprising: one or more force sensitiveprobes configured and arranged for measuring tissue deflection forcesupon application of varying forces applied on the body segment; one ormore inertial measurement units each of which include one or more ofgyroscopes, accelerometers, and magnetometers configured and arrangedfor measuring changes in tissue deflection caused by the one or moreforce sensitive probes relative to a grounded reference frame inthree-dimensional (3-D) space to create tissue deflection data; and acontroller configured to receive the tissue deflection force data andthe tissue deflection data and configured to compute segment tissueviscoelastic properties as a function of the received tissue deflectionforce data and tissue deflection data, and one or more ultrasound probespositioned at a surface of the body segment configured and arranged tomeasure ultrasound data; wherein the system is configured such that whenan operator pushes on the one or more force sensitive probes withvarying force on the body segment, the system is configured to measureat least one of tissue densities, nervous tissue transduction dynamics,blood flow, or soft tissue depths; wherein the controller is configuredto receive the tissue deflection force data and the ultrasound data andconfigured to compute the segment tissue viscoelastic properties as afunction of the tissue deflection force data and the ultrasound data;and wherein the instrument is mechanically untethered.
 2. A system fordetermining anatomical, biomechanical, and physiological properties of abody segment, including an instrument, the instrument comprising: one ormore force sensitive probes configured and arranged for measuring tissuedeflection forces upon application of varying forces applied on the bodysegment; one or more inertial measurement units each of which includeone or more of gyroscopes, accelerometers, and magnetometers configuredand arranged for performing a zero velocity update by holding theinstrument stationary at a starting point on the body segment and thenintegrating forward to calculate a trajectory in three-dimensional (3-D)space relative to the starting point to create a grounded referenceframe, and measuring changes in tissue deflection caused by the one ormore force sensitive probes relative to the grounded reference frame inthree-dimensional (3-D) space to create tissue deflection data; and acontroller configured to receive the tissue deflection force data andthe tissue deflection data and configured to compute segment tissueviscoelastic properties as a function of the received tissue deflectionforce data and tissue deflection data.
 3. The system of claim 2, whereinthe instrument is capable of wireless data transmission to a datareceiver.
 4. The system of claim 2, wherein the controller is configuredand arranged to store the measured anatomical, biomechanical, andphysiological data.
 5. The system of claim 2, wherein the instrument isa handheld probe.
 6. The system of claim 2, wherein the instrument is anarray of probes.
 7. The system of claim 2, wherein the one or more forcesensitive probes measure force using one or more force sensors that arecapacitive, resistive, piezoelectric based, strain-gauge based, orspring-potentiometer based.
 8. The system of claim 2, wherein theinstrument measures tissue stress or strain threshold where the subjectfirst experiences discomfort.
 9. The system of claim 2, wherein theinstrument comprises at least one finger probe.
 10. The system of claim9, wherein the finger probe includes a finger socket which is configuredto receive a finger of a user.
 11. The system of claim 10, wherein theforce probe is disposed on a tip of the finger socket.
 12. The system ofclaim 10, wherein the finger socket further includes a flexible armhaving a proximal end connected to the finger socket and a distal endconnected to the inertial measurement unit.
 13. The system of claim 2,further comprising: one or more ultrasound transducers to measureinternal tissue properties including tissue density and soft tissuedepth.
 14. The system of claim 13, wherein the one or more ultrasoundtransducers further measure blood flow in the body segment.
 15. Thesystem of claim 13, wherein the one or more ultrasound transducersfurther measure nervous tissue transduction dynamics.